METHOD AND APPARATUS FOR T1e-SENSITIVE INVERSION-RECOVERY IMAGING FOR EPROI

ABSTRACT

An apparatus and method for improved S/N measurements useful for electron paramagnetic resonance imaging in situ and in vivo, using high-isolation transmit/receive surface coils and temporally spaced pulses of RF energy (e.g., in some embodiments, a RF pi pulse) having an amplitude sufficient to rotate the magnetization by 180 degrees followed after varied delays, by a second RF pulse having an amplitude half that of the initial pulse to rotate the magnetization by, e.g., 90 degrees (a pi/2 pulse), to the plane orthogonal to the static field where it evolves for a short time. Then a third RF pi pulse sufficient to rotate the magnetization by, e.g., 180 degrees, forms an echo (in some embodiments, the second and third pulses are from the same signal as the first pulse but are phase shifted by 0, 90, 180, or 270 degrees to reduce signal artifact), to image human body.

RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No. 13/032,637 titled “T1-SENSITIVE INVERSION-RECOVERY-IMAGING METHOD AND APPARATUS FOR EPRI,” filed Feb. 22, 2011 (which will issue as U.S. Pat. No. 9,392,957 on Jul. 19, 2016), which claims priority to U.S. Provisional Patent Application 61/306,917 titled “HIGH-ISOLATION TRANSMIT/RECEIVE SURFACE COILS AND METHOD FOR EPRI” filed Feb. 22, 2010 by Howard J. Halpern, U.S. Provisional Patent Application 61/356,555 titled “T1-SENSITIVE INVERSION RECOVERY IMAGING APPARATUS AND METHOD FOR EPRI” filed Jun. 18, 2010 by Howard J. Halpern et al., and U.S. Provisional Patent Application 61/445,037 titled “T1-SENSITIVE INVERSION RECOVERY IMAGING METHOD AND APPARATUS FOR EPRI” filed Feb. 21, 2011 by Howard J. Halpern et al., which are all incorporated herein by reference in their entirety including their appendices.

This application is related to U.S. patent application Ser. No. 13/032,626 titled “HIGH-ISOLATION TRANSMIT/RECEIVE SURFACE COILS AND METHOD FOR EPRI” filed Feb. 22, 2011 by Howard J. Halpern (which issued as U.S. Pat. No. 8,664,955 on Mar. 4, 2014), which is incorporated herein by reference in its entirety.

FIELD OF THE INVENTION

This invention relates to the field of medical imaging and modeling, and more specifically to a method and apparatus for improved signal-to-noise (S/N) measurements useful for electron paramagnetic resonance imaging (EPRI), in situ and in vivo, using a series of transmitted pulse sets of variously temporally spaced-apart and phase-shifted pulses of transmitted radio frequency (RF) energy that disturb the electron-spin alignment, and then receive the much-weaker resulting RF signal as the electron spins realign to the static magnetic field. The transmitted signal is used in a set of static magnetic-field coils (which generate the electron-spin-aligning constant and gradient magnetic fields) and high-isolation transmit/receive surface, volume, or surface-volume coils that are configured to reduce the reception of the transmitted RF pulses by the receive coils. The received signal is then quadrature decoded (to digitized I values and Q values) using a reference RF signal. In some embodiments, a triplet set of pulses (e.g., in some embodiments, each measurement uses a triplet pulse set of three RF pulses having a selected pulse-to-pulse-to-pulse temporal spacing, a selected set of pulse amplitudes, and various selected phase-shift amounts relative to a reference RF signal, for example:

(1) a first RF “pi” pulse (e.g., about 35 ns of RF (=about 9 cycles of a signal that is about 250 MHz that cause the electron spins to rotate by 180 degrees (pi radians) relative to their orientation in the reference RF signal) of a given “full magnitude” followed, after a first predetermined time delay by (2) a second pulse of about 9 cycles of substantially the same frequency but having half the magnitude (one-quarter the power) and a shifted phase (by one of four different amounts: about 0 degrees, about 90 degrees, about 180 degrees, or about 270 degrees), relative to the cycles of the first pulse (in some embodiments, the cycles of the second pulse are obtained from the same reference RF signal source as those of the first pulse but are phase shifted by the selected amount for that set of three pulses), (the second pulse causes the directions of the electron spins to rotate 90 degrees (pi/2 radians)) and then followed, after a second predetermined time delay, by (3) a third pulse of about nine (9) cycles of substantially the same frequency but having a full magnitude (the same magnitude as the first pulse) and another shifted phase (by one of four different amounts: about 0 degrees, about 90 degrees, about 180 degrees, or about 270 degrees), relative to the cycles of the first pulse (in some embodiments, the approximately nine cycles of the third pulse are also obtained from the same signal source as those of the first pulse but are phase shifted by the selected amount and have the same full magnitude). The second pulse causes the directions of the electron spins to rotate 180 degrees (pi radians), which causes a spin echo of the T₁ relaxation signal. The T₁ spin echo is temporally separated from the third excitation pulse enough to allow better sensing (better signal-to-noise ratio (SNR)). In some embodiments, a plurality of series of pulses are used, wherein each series has sixteen (16) triplet sets of pulses, wherein each triplet set uses one of four different phase-shift amounts for the second pulse and one of four different phase-shift amounts for the third pulse, and each series uses one of a plurality of predetermined first time delays and one of a plurality of predetermined second time delays. By using pulses that sense the T₁ signal rather than the T₂ signal, the present invention provides, in some embodiments, improved micro-environmental images that are representative of particular internal structures in the human body and spatially resolved images of tissue/cell protein signals responding to conditions (such as hypoxia) that show the temporal sequence of certain biological processes, and, in some embodiments, that distinguish malignant tissue from healthy tissue.

BACKGROUND OF THE INVENTION

Cells activate protein signaling in response to crucial environmental conditions. Among the best studied is the cellular response to chronically low levels of oxygen, hypoxia. Cells respond to hypoxia by increasing hypoxia inducible factor 1α (HIF1α), a signaling peptide which is the master regulator of hypoxic response. HIF1α promotes genes and their protein products, orchestrating cell, tissue, and organism hypoxic response such as new vessel formation and increase in red cell volume.

U.S. Pat. No. 6,977,502 to David Hertz issued Dec. 20, 2005 titled “Configurable matrix receiver for MRI” is incorporated herein by reference. Hertz describes a configurable matrix receiver having a plurality of antennas that detect one or more signals. The antennas are coupled to a configurable matrix comprising a plurality of amplifiers, one or more switches that selectively couple the amplifiers in series fashion, and one or more analog-to-digital converters (ADCs) that convert the output signals generated by the amplifiers to digital form. For example, a matrix that includes a first amplifier having a first input and a first output, and a second amplifier having a second input and a second output, a switch to couple the first output of the first amplifier to the second input of the second amplifier, a first ADC coupled to the first output of the first amplifier, and a second ADC coupled to the second output of the second amplifier. In one embodiment, the signals detected by the antennas include magnetic resonance (MR) signals.

United States Patent Application Publication 2008/0084210 by Vaughan et al. published Apr. 10, 2008 titled “Multi-Current Elements for Magnetic Resonance Radio Frequency Coils” is incorporated herein by reference. Vaughan et al. disclose a current unit having two or more current paths allows control of magnitude, phase, time, frequency and position of each of element in a radio frequency coil. For each current element, the current can be adjusted as to a phase angle, frequency and magnitude. Multiple current paths of a current unit can be used for targeting multiple spatial domains or strategic combinations of the fields generated/detected by combination of elements for targeting a single domain in magnitude, phase, time, space and frequency.

United States Patent Application Publication 2008/0129298 by Vaughan et al. published Jun. 5, 2008 titled “High field magnetic resonance” is incorporated herein by reference. Vaughan et al. disclose, among other things, multi-channel magnetic resonance using a TEM coil.

An article co-authored by one inventor of the present invention is titled “Imaging radio frequency electron-spin-resonance spectrometer with high resolution and sensitivity for in vivo measurements” by Howard Halpern et al., Rev. Sci. Instrum. 60(6), June 1989, {40. Halpern, 1989 #89} was attached as Appendix B to U.S. Provisional Patent Application 61/306,917 titled “HIGH-ISOLATION TRANSMIT/RECEIVE SURFACE COILS AND METHOD FOR EPRI” filed Feb. 22, 2010 by Howard J. Halpern, which is incorporated herein by reference. Halpern et al. describe a radio frequency (RF) electron-spin-resonance spectrometer with high molar sensitivity and resolution. 250-MHz RF is chosen to obtain good penetration in animal tissue and large aqueous samples.

Another article co-authored by inventors of the present invention is titled “A Versatile High Speed 250-MHz Pulse Imager for Biomedical Applications” by Boris Epel, Sundramoorthy, S. V., Mailer, C. & Halpern, H. J. at the Center for EPR Imaging In Vivo Physiology, Department of Radiation and Cellular Oncology, University of Chicago, Chicago, Ill. 60637 (Concepts Magn. Reson. Part B (Magn. Reson. Engineering) 33B: 163-176, 2008) {46. Epel, 2008 #2200} was attached as Appendix A to U.S. Provisional Patent Application 61/306,917 titled “HIGH-ISOLATION TRANSMIT/RECEIVE SURFACE COILS AND METHOD FOR EPRI” filed Feb. 22, 2010 by Howard J. Halpern, which is also incorporated herein by reference. Epel et al. describe a versatile 250-MHz pulse electron paramagnetic resonance (EPR) instrument for imaging of small animals. Flexible design of the imager hardware and software makes it possible to use virtually any pulse EPR imaging modality. A fast pulse generation and data acquisition system based on general purpose PCI boards performs measurements with minimal additional delays. Careful design of receiver protection circuitry allowed those authors to achieve very high sensitivity of the instrument. In this article, they demonstrate the ability of the instrument to obtain three-dimensional (3D) images using the electron-spin echo (ESE) and single-point imaging (SPI) methods. In a phantom that contains a 1 mM solution of narrow line (16 μT, peak-to-peak) paramagnetic spin probe, their device achieved an acquisition time of 32 s per image with a fast 3D ESE imaging protocol. Using an 18-min 3D phase relaxation (T_(2e)) ESE imaging protocol in a homogeneous sample, a spatial resolution of 1.4 mm and a standard deviation of T_(2e) of 8.5% were achieved. When applied to in vivo imaging this precision of T_(2e) determination would be equivalent to 2 Torr resolution of oxygen partial pressure in animal tissues.

U.S. Pat. No. 4,812,763 to Schmalbein issued Mar. 14, 1989 titled “Electron spin resonance spectrometer”, and is incorporated herein by reference. Schmalbein describes an electron spin resonance spectrometer that includes a resonator containing a sample and arranged in a magnetic field of constant strength and high homogeneity. A microwave bridge can be supplied with microwave energy in the form of an intermittent signal. Measuring signals emitted by the resonator are supplied to a detector and a signal evaluation stage. A line provided between a microwave source and the microwave bridge is subdivided into parallel pulse-shaping channels, one of them containing a phase shifter, an attenuator and a switch for the signal passing through the pulse-shaping channels. In order to be able to set, if possible, an unlimited plurality of pulse sequences for experiments of all kinds, the pulse-shaping channels are supplied in equal proportions from the line by means of a divider. All pulse-shaping channels are provided with a phase shifter and an attenuator. The pulse-shaping channels are re-united by means of a combiner arranged before the input of a common microwave power amplifier.

U.S. Pat. No. 6,639,406 to Boskamp, et al. issued Oct. 28, 2003 titled “Method and apparatus for decoupling quadrature phased array coils”, and is incorporated herein by reference. Boskamp, et al. describe a method and apparatus for combining the respective readout signals for a loop and butterfly coil pair of a quadrature phased array used for magnetic resonance imaging. The technique used to combine the signals introduces a 180-degree phase shift, or multiple thereof, to the loop coil signal, thereby allowing the loop coil signal to be decoupled from other loop coil signals by a low-input-impedance preamplifier in series with the signal. This patent describes a surface coil that is applied to one surface of the body part being examined.

U.S. Pat. No. 7,659,719 to Vaughan, et al. issued Feb. 9, 2010 titled “Cavity resonator for magnetic resonance systems”, and is incorporated herein by reference. Vaughan, et al. describe a magnetic resonance apparatus that includes one or more of the following features: (a) a coil having at least two sections, (b) the at least two sections having a resonant circuit, (c) the at least two sections being reactively coupled or decoupled, (d) the at least two sections being separable, (e) the coil having openings allowing a subject to see or hear and to be accessed through the coil, (f) a cushioned head restraint, and (g) a subject input/output device providing visual data to the subject, the input/output device being selected from the group consisting of mirrors, prisms, video monitors, LCD devices, and optical motion trackers. This patent describes a volume head coil that surrounds a human head.

U.S. Pat. No. 5,706,805 Swartz, et al. issued Jan. 13, 1998 titled “Apparatus and methodology for determining oxygen tension in biological systems”, and is incorporated herein by reference. Swartz, et al. describe apparatus and methods for measuring oxygen tensions in biological systems utilizing physiologically acceptable paramagnetic material, such as India ink or carbon black, and electron paramagnetic resonance (EPR) oximetry. India ink is introduced to the biological system and exposed to a magnetic field and an electromagnetic field in the 1-2 GHz range. The EPR spectrum is then measured at the biological system to determine oxygen concentration. The EPR spectrum is determined by an EPR spectrometer that adjusts the resonator to a single resonator frequency to compensate for movements of the biological system, such as a human or animal. The biological system can also include other in vivo tissues, cells, and cell cultures to directly measure pO₂ non-destructively. The paramagnetic material can be used non-invasively or invasively depending on the goals of the pO₂ measurement. A detecting inductive element, as part of the EPR spectrometer resonator, is adapted relative to the measurement particularities.

U.S. Pat. No. 5,865,746 to Murugesan, et al. issued Feb. 2, 1999 titled “In vivo imaging and oxymetry by pulsed radiofrequency paramagnetic resonance”, and is incorporated herein by reference. Murugesan et al. describe a system for performing pulsed RF FT EPR spectroscopy and imaging includes an ultra-fast excitation subsystem and an ultra-fast data acquisition subsystem. Additionally, method for measuring and imaging in vivo oxygen and free radicals or for performing RF FT EPR spectroscopy utilizes short RF excitations pulses and ultra-fast sampling, digitizing, and summing steps.

U.S. Pat. No. 4,280,096 to Karthe, et al. issued Jul. 21, 1981 titled “Spectrometer for measuring spatial distributions of paramagnetic centers in solid bodies”, and is incorporated herein by reference. Karthe, et al. describe a spectrometer in which gradient coils are provided in order to create an inhomogeneous magnetic field for use in analyzing individual regions within the sample under examination. The gradient coils and the modulating coils are operated by discrete pulses, rather than continuously. A keying unit coordinates the interaction of the various components of the spectrometer in order to monitor resonance of the sample under examination while such pulses occur.

U.S. Pat. No. 5,828,216 to Tschudin, et al. issued Oct. 27, 1998 titled “Gated RF preamplifier for use in pulsed radiofrequency electron paramagnetic resonance and MRI”, and is incorporated herein by reference. Tschudin et al. describe a gated RF preamplifier used in system for performing pulsed RF FT EPR spectroscopy and imaging or MRI. The RF preamplifier does not overload during a transmit cycle so that recovery is very fast to provide for ultra-fast data acquisition in an ultra-fast excitation subsystem. The preamplifier includes multiple low-gain amplification stages with high-speed RF gates inserted between stages that are switched off to prevent each stage from overloading during the transmit cycle.

U.S. Pat. No. 4,714,886 to one of the present inventors, Howard Halpern, issued Dec. 22, 1987 titled “Magnetic resonance analysis of substances in samples that include dissipative material”, and is incorporated herein by reference. U.S. Pat. No. 4,714,886 describes magnetic resonance images of the distribution of a substance within a sample that are obtained by splaying a pair of magnetic field generating coils relative to each other to generate a magnetic field gradient along an axis of the sample. In other aspects, electron spin resonance data is derived from animal tissue, or images are derived from a sample that includes dissipative material, using a radio frequency signal of sufficiently low frequency.

There is a need for an improved apparatus and method of electron-spin-resonance spectrometry and/or imaging to non-invasively provide images and/or other signal measurements representative of particular internal structures and processes in the human body, and to be able to distinguish malignant tissue from healthy tissue.

SUMMARY OF THE INVENTION

The present invention provides a method and apparatus for improved signal-to-noise (S/N) measurements useful for electron paramagnetic resonance imaging (EPRI), in situ and in vivo, using high-isolation transmit/receive surface coils and a series of pulse sets of variously temporally spaced-apart and phase-shifted pulses of RF energy. In some embodiments, a plurality of measurements are taken and recorded, wherein each measurement is based on establishing a static (DC) magnetic field on an animal tissue (e.g., a magnetic field having a substantially fixed direction and a gradient field strength in a section of volume of tissue of a living human), then transmitting a set of RF pulses that includes a first pulse, a first delay, a second pulse, a second delay and a third pulse, and then receiving, processing, and storing the resultant RF signals from the tissue sample. By varying the strength of the DC magnetic field (e.g., by generating the magnetic-field gradient that increases the field strength in some areas of the tissue being measured and/or decreases the field strength in other areas of the tissue being measured) and varying the direction of the gradient relative to the tissue being measured (e.g., by electro-magnetically changing the direction of the gradient, or by physically moving the patient relative to the gradient), varying the durations of the first and second delays, and varying the amounts of phase shifts of the second and third pulses relative to the phase of the first pulse, and storing that data along with data based on the RF signal received from the sample, an image (e.g., two-dimensional (2D) sections in various orientations, or a three-dimensional (3D) image of a volume) of the section or volume of tissue can be derived by computer calculations and displayed. For example, in some embodiments, each measurement uses a pulse set of three RF pulses having a selected temporal spacing, selected amplitudes, and various phase-shift amounts, for example: a first RF pulse (e.g., in some embodiments, a pulse of about 35-ns duration, which equals about 9 cycles of about 250 MHz cycles) followed after a first variable delay by a second pulse (e.g., of about 9 cycles of substantially the same frequency but having half the amplitude and a shifted phase relative to the cycles of the first pulse; in some embodiments, the cycles of the second pulse are obtained from the same signal source as those of the first pulse but are either not phase shifted (which is equivalent to phase shifted by 0 degrees), phase shifted by 90 degrees, phase shifted by 180 degrees, or phase shifted by 270 degrees, then followed, after a second variable delay, by a third pulse of about 9 cycles of substantially the same frequency but having the same amplitude as the first pulse and a variously shifted phase relative to the cycles of the first pulse. In some embodiments, the first pulse has a magnitude that flips the electron-spin directions (the direction of magnetization) by 180 degrees and is thus called a “pi pulse”; the second pulse has a magnitude that flips the electron-spin directions (the direction of magnetization) by 90 degrees and is thus called a “pi-over-two pulse” (or “pi/2 pulse”); and the third pulse has a magnitude that again flips the electron-spin directions (the direction of magnetization) by 180 degrees and is thus called a “pi pulse”. In some embodiments, these triplet sets of excitation RF pulses are configured to result in signals that represent the T₁ relaxation of electron spin (in contrast to the T₂ relaxation of electron spin) in order to be sensitive to the oxygen content in the tissue being imaged. In some such embodiments, the four possible phase-shift amounts of the second pulse and the four possible phase-shift amounts of the third pulse provide sixteen combinations of phase shift amounts for each set of the first delay (between the first and second pulses) and second delay (between the second and third pulses). In some embodiments, this provides improved SNR in micro-environmental images that are representative of particular internal structures in the human body and spatially resolved images of tissue/cell protein signals responding to conditions (such as hypoxia) that show the temporal sequence of certain biological processes, and, in some embodiments, that distinguish malignant tissue from healthy tissue.

In some embodiments, the durations of the three pulses in a triplet set are kept constant and the frequency of the carrier in each pulse are kept constant in order to maintain the same spectral content (i.e., the Fourier transform of the pulses yields the same spectrum of frequencies and relative strengths for those frequencies) for every pulse, while the total strength of some of the pulses is varied to obtain different amounts of rotation of the electron spins and/or magnetic moment of the reporter molecules.

In some embodiments, the present invention further includes medical procedures, animal models, and biological agents (such as viral “Trojan Horse” constructs or other vectors) that facilitate the obtaining of images that distinguish different types of tissues or healthy tissues from malignant or infected tissues, that show various spatially and temporally resolved signaling, regulation, promotion and responses of, for example, signaling peptides, protein products.

BRIEF DESCRIPTION OF THE DRAWINGS

The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.

FIG. 1A1 shows a schematic representation of an inversion-recovery pulse sequence 101 useful for measuring T₁, according to some embodiments of the present invention.

FIG. 1B1 shows a schematic representation of a conventional saturation-recovery-with-echo-detection pulse sequence 102 useful for measuring T₁.

FIG. 1C1 shows a schematic representation of a 2 pESE pulse sequence 103 useful for measuring T₂, according to some embodiments of the present invention.

FIG. 1A2 shows another schematic representation of inversion-recovery pulse sequence 101 useful for measuring T₁, according to some embodiments of the present invention.

FIG. 1B2 shows another schematic representation of conventional saturation-recovery-with-echo-detection pulse sequence 102 useful for measuring T₁.

FIG. 1C2 shows another schematic representation of 2 pESE pulse sequence 103 useful for measuring T₂, according to some embodiments of the present invention.

FIG. 1D is a perspective schematic representation of a gradient-field set 104 of electromagnet coils useful for measuring T₁, according to some embodiments of the present invention.

FIG. 2 is a graph 201 of spectroscopic measurements of the effective spin-packet linewidth versus spin-probe concentration.

FIG. 3A and FIG. 3B present 1/T₂ and 1/T₁ images, respectively, obtained from a mouse tumor, presented on the same drawing sheet for comparison.

FIG. 3C is an enlarged view of the screenshot of the 1/T₂ images of FIG. 3A.

FIG. 3D is an enlarged view of the screenshot of the 1/T₁ images of FIG. 3B.

FIG. 4A shows an SFR ESE pulse sequence, wherein the repetition rate is varied.

FIG. 4B shows an SFR SPI pulse sequence, wherein the repetition rate is varied.

FIG. 4C shows a saturation-recovery-with-echo-detection (SR) pulse sequence, wherein the delay time T is varied.

FIG. 4D1 shows an IRESE pulse sequence using a pi/2-pi echo-detecting pulse pair, wherein the delay time T is varied.

FIG. 4D2 shows an IRESE pulse sequence with a pi/2-pi/2 echo-detecting pulse pair, wherein the delay time T is varied.

FIG. 4E shows an IRSPI pulse sequence, wherein the delay time T is varied.

FIG. 4F shows an SE pulse sequence, wherein the delay time T is varied.

FIG. 4G shows a two-pulse electron-spin echo (2 pESE) sequence measuring T₁ by varying repetition time.

FIG. 5A shows a selected slice T_(1e) image and histogram obtained using SFR ESE.

FIG. 5B shows a selected slice T_(1e) image and histogram obtained using SFR SPI.

FIG. 5C shows a selected slice T_(1e) image and histogram obtained using IRESE.

FIG. 5D shows a selected slice T_(1e) image and histogram obtained using IRSPI.

FIG. 5E shows a selected slice T_(1e) image and histogram obtained using SE.

FIG. 5F shows a selected slice T_(2e) image and histogram obtained using 2 pESE.

FIG. 6A is a T_(2e) image obtained using a 2 pESE sequence.

FIG. 6B is a T_(1e) image obtained using an IRESE sequence.

FIG. 7 is an EPR oxygen image of two planes of a mouse leg bearing an FSa fibrosarcoma.

DESCRIPTION OF PREFERRED EMBODIMENTS

Although the following detailed description contains many specifics for the purpose of illustration, a person of ordinary skill in the art will appreciate that many variations and alterations to the following details are within the scope of the invention. Accordingly, the following preferred embodiments of the invention are set forth without any loss of generality to, and without imposing limitations upon, the claimed invention. Further, in the following detailed description of the preferred embodiments, reference is made to the accompanying drawings that form a part hereof, and in which are shown by way of illustration specific embodiments in which the invention may be practiced. It is understood that other embodiments may be utilized and structural changes may be made without departing from the scope of the present invention.

The leading digit(s) of reference numbers appearing in the Figures generally corresponds to the Figure number in which that component is first introduced, such that the same reference number is used throughout to refer to an identical component which appears in multiple Figures. Signals and connections may be referred to by the same reference number or label, and the actual meaning will be clear from its use in the context of the description.

Section #1: T₁ Imaging

Virtually all pulse-EPR imaging in animal specimens has used sequences sensitive to transverse relaxation times of the electron spin probe, T₂. {110. Matsumoto, 2006}, {46. Epel, 2008 #2200}. T₂ is a measure of phase coherence that takes place as a result of quantum-mechanical spin flips caused by multiple interactions with the environment. As has been demonstrated in nuclear-magnetic-resonance imaging and spectroscopy, and water-proton MRI, T₂ relaxation is sensitive not only to the loss of energy by the spin system to the environment (spin lattice) but multiple other relaxation processes that can affect the phase coherence of the magnetization spins as well as the loss of energy. Among these relaxation processes are interactions, in animal tissue, with oxygen, a di-radical with two unpaired electrons in its outermost orbital. This both dephases the spins of the reporter spin system but results in loss of energy from the reporter spin system.

One of the major confounding and blurring effects of using T₂ imaging for oxygen measurement is the confounding spin dephasing by one molecule of the reporter system by another reporter molecule. These cannot be distinguished in a T₂ measurement of oxygen dephasing of the reporter spin system.

T₂ measurements are also known as transverse-spin-relaxation time (or sometimes spin-spin relaxation time) measurements. T₁ measurements are also known as longitudinal-spin-relaxation-time measurements. In principle, T₁ only measures the loss of energy by the reporter spin system to the environment, also known as the lattice, and thus T₁ is also called the spin-lattice relaxation time (SLR time), while T₂ is also called spin-spin relaxation time. This requires that the mechanism(s) that can relax the longitudinal spins are much more restricted than those that relax the transverse spin component. T₁-spin relaxation is thereby a more specific process. An example, in principle, is that when energy is transferred from one reporter spin molecule to another, the reporter spin system loses no energy. There is no relaxation. Thus, the T₁ measurement should be far less sensitive to self interaction than the T₂ measurement.

The present invention has obtained the first T₁-based tissue-oxygen image using T₁-sensitive inversion recovery using spin-echo detection of the spin recovery. This gives longer relaxation times, which will provide an even more direct sensitivity to pO₂.

Two modalities for determination of T₁ are commonly used in pulse EPR: saturation and inversion recovery. In the first case, presented in the FIG. 1B1, the EPR transition is saturated with a long (t>>T₁) RF pulse and recovery of EPR signal is detected as a function of separation between this pulse and detection sequence (two pulse echo in this case). For inversion recovery (FIG. 1A1), a π-pulse (pi pulse) is used to invert the magnetization. We have implemented the inversion-recovery method since it has a twice larger effect and does not require additional hardware. In some embodiments, for projection generation, which acts as a “read out” of the T₁ information, the present invention uses the electron-spin echo (ESE)-detection sequence.

In an EPR system, the magnetic moment will generally align with the H₀ static magnetic field (conventionally and as used herein the direction of H₀ is designated as the z direction), but will precess at an angular frequency ω that is proportional to H₀, and there exists a particular angular frequency ω_(R) that is resonant for a particular magnitude of H₀. If an alternating magnetic field H₁ cos(ωt) (designated H₁ for simplicity, conventionally and as used herein the direction of H₁ is designated as the x direction) is applied orthogonal to the H₀ static magnetic field, where ω is at the resonant angular frequency ω_(R), a magnetic moment that was parallel to the H₀ static magnetic field will precess in the y-z plane; that is, the magnetic moment will precess but always remain perpendicular to the H₁ (x direction) alternating field, and thus will periodically be pointed in a direction opposite H₀. If a wave train (a radio-frequency (RF) pulse) were applied at a magnitude A and for a pulse duration t_(w) the magnetic moment will precess through an angle θ=AH₁t_(w). If an RF pulse of a selected magnitude A and pulse duration t_(w) is applied such that θ=π (i.e., θ=180 degrees), the pulse will invert the magnetic moment and such a pulse is called a π pulse (also called a 180-degree pulse). Further, if an RF pulse of a different selected magnitude A′ and pulse duration t_(w) is applied such that θ=π/2 (i.e., θ=90 degrees), the pulse will turn the magnetic moment from the z direction to the y direction and such a pulse is called an/2 pulse (also called a 90-degree pulse). Note that either or both the magnitude A and pulse duration t_(w) may be varied to achieve a given desired rotation of the magnetic moment. (See Charles P. Slichter “Principles of Magnetic Resonance, Third Enlarged and Updated Edition”, Springer Berlin-Heidleberg, 1996, pp 20-24.) Other pulses of selected magnitudes A and pulse durations t_(w) may be used to achieve other amounts of rotation of the magnetic moment, such as a 2π/3 pulse (also called a 120-degree pulse).

FIG. 1A1 shows a schematic representation of an inversion-recovery pulse sequence 101 that includes a first π pulse (also called a pi pulse, inversion pulse or 180-degree pulse) 110, a first time delay “T” 111 after first π pulse (pi pulse) 110, then a second π/2 pulse (pi/2 pulse or 90-degree pulse) 112, a second time delay “τ” 113 after second π/2 pulse (pi/2 pulse) 112; then a third π pulse (pi pulse or 180-degree pulse) 114, a third time delay “τ” 115 after third r pulse (pi pulse) 114; then a readout-stage measurement 116 of the resulting spin echo. This pulse sequence is referred to herein as an inversion recovery sequence with electron-spin echo detection, or IRESE. The dashed line 127 shows signal amplitude (arbitrary units; not to scale) of the spin echo as a function of time. In both cases (FIG. 1A1 and FIG. 1B1), the detection sequence consists of two ESE pulses: a π/2 pulse (pi/2 pulse or 90-degree pulse) and a π pulse (pi pulse or 180-degree pulse). Note that the durations of the pulses as shown in FIG. 1A1, FIG. 1B1, and FIG. 1C1 are not to scale with the times between pulses or times between pulse sequences. In some embodiments, the pulse durations are chosen as 35 ns (nanoseconds) each (substantially square pulses of about 9 cycles of a 250 MHz carrier wave), while the time τ 113 between the π/2 pulse 112 and the π pulse 114 is chosen as τ=630 ns, and T 111 is varied for a given set of signal acquisitions at eight values, denoted as variable delay or VD in the IRESE section of Table 1 below, logarithmically spaced between 0.5 μs (500 ns) and 16 μs (16,000 ns). Further, the magnitudes of the readout signals 116, 126, and 136 are not to scale with the magnitude of the excitation pulses 110, 112, 114, 120, 122, 124, 132, or 134.

Pulse 112 and pulse 114 are also referred to as “readout pulses”. In some other embodiments, rather than using a π/2 pulse (90-degree pulse) and a π pulse (180-degree pulse) as the readout pulses, two pulses, each being a 2π/3 pulse (120-degree pulse), are used instead as the readout pulses for an inversion-recovery measurement of T₁. The original spin echo measurements of Erwin Hahn, which were generated by two π/2 pulses, could also be used for the readout of the magnetization after an initial inversion pulse

FIG. 1B1 shows a schematic representation of a saturation-recovery-with-echo-detection (SR) pulse sequence 102 that includes a first saturation pulse 120, a first time delay “T” 111 after first saturation pulse 120, then a second π/2 pulse (pi/2 pulse or 90-degree pulse) 122, a second time delay “τ” 123 after the second pulse 122; then a third π pulse (pi pulse or 180-degree pulse) 124, a third time delay “τ” 125 after the third pulse 124; then a readout-stage measurement 126 of the resulting spin echo. The dashed line 127 shows relative signal amplitude (arbitrary units; not to scale) of the spin echo as a function of time.

FIG. 1C1 shows a schematic representation of a 2 pESE pulse sequence 103 useful for measuring T₂, wherein pulse sequence 103 includes a first π/2 pulse 132, a first time delay “τ” 133 after the first pulse 132; then a second π pulse 134, a second time delay “τ” 135 after the second pulse 134; then a readout-stage measurement 136 of the resulting spin echo. The delay time T_(R) 139 is the time between the end of one signal-acquisition phase (e.g., of signal 136) and the start of the next first pulse 132. The dashed line 127 shows signal amplitude (arbitrary units; not to scale) of the spin echo as a function of time.

FIG. 1A2 shows another schematic representation of inversion-recovery pulse sequence 101 useful for measuring T₁, according to some embodiments of the present invention. Note that the durations of the pulses as shown in FIG. 1A2, FIG. 1B2, and FIG. 1C2 are more to scale with the times between pulses or times between pulse sequences. In some embodiments, as described above, the pulse durations are chosen as 35 ns (nanoseconds) each (substantially square pulses of about 9 cycles of a 250 MHz carrier wave), while the time τ 113 between the π/2 pulse 112 and the π pulse 114 is chosen as τ=630 ns, and T 111 is varied for a given set of signal acquisitions at eight values, denoted as VD in the IRESE section of Table 1 below, logarithmically spaced between 0.5 μs (500 ns) and 16 μs (16000 ns). Further, the magnitudes of the readout signals 116, 126, and 136 are not to scale with the magnitude of the excitation pulses 110, 112, 114, 120, 122, 124, 132, or 134.

FIG. 1B2 shows another schematic representation of conventional saturation-recovery-with-echo-detection pulse sequence 102 useful for measuring T₂, where the durations of the pulses and the times between pulses are more to scale than in FIG. 1B1.

FIG. 1C2 shows another schematic representation of 2 pESE pulse sequence 103 useful for measuring T₁, according to some embodiments of the present invention, where the durations of the pulses and the times between pulses are more to scale than in FIG. 1C1.

FIG. 1D is a perspective schematic representation of a gradient-field set 104 of electromagnet coils useful for measuring T₁, according to some embodiments of the present invention. In some embodiments, the planes of all of the rectangular coils (coil 142LT (the left-hand top coil in FIG. 1D), coil 142RT (the right-hand top coil in FIG. 1D), coil 142LB (the left-hand bottom coil in FIG. 1D), coil 142RB (the right-hand bottom coil in FIG. 1D), coil 143FT (the front top coil in FIG. 1D), coil 143BT (the back top coil in FIG. 1D), coil 143FB (the front bottom coil in FIG. 1D), coil 143BB (the back bottom coil in FIG. 1D), are parallel to one another and these planes in turn are parallel to the planes of the round coils. This is to make the main direction of the magnetic fields of each given strength in the region of interest parallel. In some embodiments, the planes of the pairs of similarly oriented square coils, for example, above the horizontal plane, are offset a little bit so that these rectangular coil pairs can be moved closer to each other without bumping each other. In some embodiments, the distances between members of each pair are reduced to make the gradient (the amount of change in magnetic field strength) more uniform.

In some embodiments, the currents in each coplanar pair move in opposite senses. For example, in the embodiment shown in FIG. 1D, the current in coil 141T is clockwise such that the differential field 151T is downward (subtracting from the main field 150 on the top side), while the current in coil 141B is counterclockwise such that the differential field 151B is up (adding to the main field 150 on the back side). By changing the magnitude of the currents in coils 141T and 141B the amount of gradient can be varied. Similarly, in the embodiment shown, the current in coils 142LT and 142LB are clockwise such that the differential field 152L is downward (subtracting from the main field 150 on the left-hand side), while the current in coils 142RT and 142RB are counterclockwise such that the differential field 152R is upward (adding to the main field 150 on the right-hand side); and the current in coils 143FT and 143FB are counterclockwise such that the differential field 153F is upward (adding to the main field 150 on the front side), while the current in coils 143BT and 143BB are clockwise such that the differential field 153B is downward (subtracting from the main field 150 on the back side). By changing the directions and the relative and absolute amounts of current in the three sets of coils, the gradient-field direction and magnitude can be changed in plus and/or minus X, Y and/or Z directions.

In some embodiments of the gradient-field coils, there are left and right coplanar pairs (142LB and 142RB) that face the corresponding coils (142LT and 142RT) above the horizontal midplane, and front and back coplanar pairs (143FB and 143BB) of coils below the horizontal midplane that face the corresponding coils (143FT and 143BT) above the horizontal midplane. In some embodiments (not shown here), for each pair of the gradient-field coils (142LB and 142RB), (142LT and 142RT), (143FB and 143BB) (143FT and 143BT) there are an outside set of two pairs and an inside set of two pairs, one set closer—the other set farther apart. The coils that face each other across the horizontal have their currents in the same direction.

FIG. 2 presents a graph 201, where plot 211 shows 1/T₁, the longitudinal relaxation rate of the magnetization described in terms of magnetic-field linewidth units to which it is proportional versus spin-probe concentration as measured using an inversion-recovery pulse sequence 101, and plot 212 shows 1/T₂, the transverse relaxation rate of the magnetization described in terms of a magnetic field linewidth to which it is proportional versus spin-probe concentration as measured using a spin echo pulse sequence decay 103.

FIG. 2 presents a graph 201 of spectroscopic measurements of a the effective spin-packet linewidth versus spin-probe concentration of deoxygenated samples of spin probe (OX063 radical methyl-tris[8-carboxy-2,2,6,6-tetrakis[2-hydroxyethyl]benzo[1,2-d:4,5-d′]bis[1,3]dithiol-4-yl]-trisodium salt, molecular weight=1,427 from GE Healthcare, Little Chalfont, Buckinghamshire, United Kingdom). In some embodiments, the solvent for the spin probe is saline. All measurements are made at the same oxygen partial pressure, 0 torr. Plot 211 shows the T₁ spin-packet linewidth versus spin-probe concentration as measured using an inversion-recovery (IRESE) pulse sequence 101 (see FIG. 1A1) according to the present invention, while plot 212 shows the T₂ spin-packet linewidth versus spin-probe concentration as measured using a saturation-recovery-with-echo-detection pulse sequence 102 (see FIG. 1B1). Note that at low spin-probe concentrations, the T₁ and T₂ measurements give approximately the same linewidth. However, as concentration of spin probe rises, the additional confounding width of the T₁ measurements is reduced, and is only about one-fifth (⅕) that of the T₂ measurements. This is a remarkable reduction in the sensitivity of linewidth (or relaxation time) to the confounding variation of the concentration of spin probe. That is, the linewidth increase (i.e., error and/or uncertainty due to spin-probe concentration increase) is about five times worse at high spin-probe concentrations if measuring T₂ using saturation-recovery-with-echo-detection pulse (SR) sequence 102 than if measuring T₁ using an inversion-recovery pulse (IRESE) sequence 101.

FIG. 3A and FIG. 3B present T₂ and T₁ images obtained from the same tumor, presented on the same drawing sheet for comparison. FIG. 3C is an enlarged view of the screenshot of FIG. 3A, and FIG. 3D is an enlarged view of the screenshot of FIG. 3B.

FIG. 3A and FIG. 3C each show, at different magnifications, the same orthogonal T₂ EPR image planes 301, 302 and 303 (labeled YX, ZY and YZ respectively) from an C57/B mouse leg bearing a syngeneic B 16 melanoma. The red contour line in each view is that of the tumor from obtained from a spatially registered MRI (a conventional nuclear magnetic resonance image). Spin-packet linewidth derived pO₂ from the whole leg are histogrammed in blue plot 304 and those only from the tumor are in red plot 305.

FIG. 3B and FIG. 3D each show, at different magnifications, the same orthogonal T₁ EPR image planes 306, 307 and 308 (labeled YX, ZY and YZ respectively) from the same C57/B mouse leg bearing the syngeneic B 16 melanoma. The red contour line in each view is that of the tumor from a conventional nuclear MRI. Spin-packet linewidth derived pO₂ from the whole leg are histogrammed in blue plot 309 and those only from the tumor are in red plot 310.

From the histograms 304, 305 of the pO₂ values derived from the T₂ images and the histograms 304, 305 of the pO₂ values derived from the T₁ images one can see a clear separation of the hypoxic tumor pO₂s in the T₁ image which is blurred in the T₂ image. This is due to the reduced confounding effect of spin probe concentration on the images.

Section #2: Further T₁ Imaging Techniques—

a. Significance—Oxygen, HIF1α, VEGF and Cancer Biology:

Regions of low pO₂-hypoxia—are characteristic of solid tumors and have long been known to increase resistance of malignant cells to radiation {1. Hall, 2000 #1012; 2. Gatenby, 1988 #21; 3. Brizel, 1996 #695; 4. Brizel, 1999 #1121; 5. Brizel, 1996 #1124; 6. Brizel, 1997 #1123; 7. Hockel, 1996 #1111} and can be exploited for cancer therapy. 18. Shibata, 2002 #1657.1 Hypoxia selects for a mutagenic, carcinogenetic, and aggressive malignant phenotype. 19. Graeber, 1996 #942.1 Oxygen status is so important in tissue homeostasis that the absence of oxygen, hypoxia, is the causative element in an entire regulatory peptide signaling cascade. Hypoxia inducible factor 1α, HIF1α, is the master regulator of the response of the cell response to hypoxia, initiating the signal cascade. {10. Semenza, 1998 #1693.} The cascade generates compensatory responses to hypoxia at the cellular level, an intracrine response e.g., apoptosis, {11. Carmeliet, 1998 #1132}, local vascular response through increased production of Vascular Endothelial Growth Factor (VEGF), a paracrine response {124. Carmeliet, 2000 #1131; 10. Semenza, 1998 #1693}, and a general organism response, e.g., erythrocyte production through the increase in production of erythropoietin, an endocrine response {10. Semenza, 1998 #1693.}

The anti-cancer success of anti-VEGF therapy {128. Sandler, 2007 #2192} as well as in vitro work, e.g., {126. Li, 2007 #2214; 125. Forsythe, 1996 #2109} argues that in vivo, in situ, the signal peptide response to hypoxic environment is different in normal and malignant cells and tissues. The present invention, for the first time, uses electron paramagnetic resonance (EPR) imaging (EPRI) to demonstrate this non-invasively in animals with possible extension to humans. Most tests of cell signaling in vivo model subjected the whole animal to an environmental change and then imaged the local response ex vivo. {127. Picchio, 2008 #2103.} There is virtually no literature dealing with the heterogeneity in normal tissue and solid tumors. Although in vitro work has established HIF1α signaling, this has profound implications for therapeutic strategy. For example, it may explain failure of oxygen manipulation to enhance tumor therapy with radiation. {129. Suit, 1984 #1584}

A major goal of the Center for EPR Imaging In Vivo Physiology at the University of Chicago is obtaining uniquely high-resolution quantitative images of tissue and tumor pO₂. The present invention provides an entirely new means of molecular imaging of the peptide signal response to hypoxia, using EPR, registered with pO₂ images. This technique can be extended to a vast array of peptide-signaling processes. An example of this is the imaging of VEGF, a HIF1α cascade signal response to create new vessels. Combined, co-localized images of pO₂ and peptide-signaling response will produce a quantified, localized relationship between the extent of hypoxia and the cell/tissue signal response to hypoxia in vivo, in situ, which we hypothesize is different in malignant and normal tissue. We anticipate the eventual extension of EPR imaging technology to humans. pO₂ stimulus images registered with peptide signal response will show individual variations in local stimulus-response to guide individual therapy. The cell signaling technology described here will impact the study of human health and disease.

FIG. 8 is an EPR oxygen image of two planes of a mouse leg bearing an FSa fibrosarcoma. Colorbars show pO₂ in mm/hg (torr). Numbers on planes are mm. Resolution: 1 mm spatial, 3 torr pO₂. Tumor is not distinguished in this image but separately defined using a registered T₂ MRI indicating large central oxygen gradients in the tumor.

Molecular Imaging Shows Heterogeneity of Tumor/Tissue Condition and Signal Response:

Although cell signaling discoveries have provided unique insight into modes by which cells communicate with cells in their environment, {14. Alberts, 2008 #2096} these studies of isolated cells in artificial homogeneous environments contrast with the enormous heterogeneity of a living animal as shown in FIG. 4 and {15. Fischbach, 2009 #2089}. The interaction of anatomy and signaling molecules through vascular bed structure, target organ distance, size and location can affect signaling. Organ- or tissue-dependent modulation of the signaling can provide another layer of control that needs to be understood to fully comprehend the physiology of signaling. A crucial reason to image cell signaling is this variation with position of cellular environments, shown in FIG. 1. Registering images of a quantified environment characteristic like pO₂ with images of the peptide response allows the development of models stimulus and response in a native environment.

Reporter Protein (RP)/Molecular Beacon (MB) Imaging Technologies:

The reporter gene LacZ has been a major tool used to dissect transcription induction using optical and fluorescence molecular beacon (MB) detection. {16. Alam, 1990 #1662.} The original such technology, LacZ, the bacterial gene encoding β-galactosidase (reporter) turns the indole linked sugar X-Gal blue, an optical MB. {Holt, 1958 #1663.} By coupling the β-galactosidase gene to a gene of interest, gene expression is directly seen taking place in blue cells. Many other such technologies have followed. {16. Alam, 1990 #1662; 18. Chalfie, 1994 #1664; 19. Weissleder, 2003 #1694; 20. McCaffrey, 2003 #1692; 21. Blasberg, 2003 #1681; 22. Massoud, 2003 #1687; 23. Herschman, 2003 #1684}, producing chromophore or fluorescent MBs in cells producing transcriptionally coupled gene products that can be detected and imaged in vivo. The work proposed in this grant uses this basic technique, modified to turn on an EPR MB in response to hypoxia simultaneously imaged in vivo, obviating the problems with radionuclide, optical or MRI techniques.

Molecular Imaging Techniques Other than EPR do not Easily Allow Imaging of Animal Environment Condition and Cell Signal Response:

Optical images and radionuclide imaging dominate molecular imaging. {24. Dothager, 2009 #2227.} Optical techniques (e.g., www.xenogen.com/prodimag1.html) use reporter genes that can be engineered into transgenic mice {25. Zhang, 2001 #1695} or into implanted tumor cells in mice {26. Adams, 2002 #1680} are surface weighted because of the rapid non-resonant absorption of optical frequency light by tissue. This makes it difficult to quantify image signal intensity, linewidths or relaxation times of depth greater than a few mm. {27. Kirkpatrick, 2004 #1691.} Other than in artificial systems such as window chamber {28. Dewhirst, 1996 #1775}, quantified relationship between stimulus such as micro-environmental oxygen and peptide signal response is difficult.

Detection of radiotracer with positron emission tomography (PET) avoids problems with depth sensitivity {29. Schober, 2009 #2099; 30. Sun, 2001 #1374; 31. Blasberg, 2003 #1665} and is extremely flexible. The advantage of radiotracer reporters is that it can be immediately translated to human studies. However, the major problems with PET imaging is its limited resolution in space (˜2 mm) and in time and distinguishing signal from the environmental stimulus reporter from the peptide signal response reporter. For radionuclide studies, hypoxia is defined as the reductive retention of nitro-imidazole {33. Raleigh, 1992 #765; 34. Evans, 1996 #931} or ATSM copper chelates. {35. Lewis, 1999 #1371.} Hypoxic signaling via HIF1α might be imaged as is proposed in this grant with vectors containing hypoxia responsive elements (HREs) that bind HIF1α promoting production of, e.g., a thymidine kinase RP that would cause hypoxic cells to retain radioactive thymidine (the molecular beacon (MB)). This has severe limitations:

Firstly, it is difficult to distinguish the signal from the compound signaling hypoxia from the thymidine retained through phosphorylation, signaling hypoxic response. EPR allows spectrally distinct hypoxia images and the peptide signal response images.

Secondly, the EPROI is quantitative while the reductive retention image is qualitative. Radionuclide images depend heavily on access of the radionuclide to the location where oxygen is measured, and other aspects of local tissue reductive capability, i.e., P450 reductase, xanthine oxidase, etc. activity. {36. Melo, 2000 #2229.} For EPROI, as long as some spin probe reaches the location, the oxygen measurement depends only weakly on the signal amplitude. Rather it depends on the signal relaxation time or line width.

Molecular imaging with MRI creates contrast with the RP requiring an extremely large molecular signal {37. Louie, 2000 #1395; 38. Weissleder, 2000 #1673} because the technology introduces contrast in a very high signal background. Unlike MRI, the EPR technology activates a “beacon in the dark”. In addition, MRI images provide poor pO₂ sensitivity.

EPR Oxygen Images Use Very Low Magnetic Fields and are Specific and Sensitive to pO₂:

EPR images are obtained at excitation RF frequencies of very-high-field (6-7 T) whole-body MRIs {39. Vaughan, 2009 #2230} but because the magnetic moment of the electron is 658 times that of the water proton, magnetic fields are 1/658 that of MRI. EPRI uses a low field with inexpensive magnet systems of about 90 Gauss=9 milliTesla (mT) at 250 MHz frequency. {40. Halpern, 1989 #89; 41. Halpern, 1991 #899.} This is a low-cost technology not requiring superconducting magnets, although standard-field MRI is, in some embodiments, used to identify tumor. EPR spectral linewidths of certain carbon-centered spin probes, trityls, are specific and sensitive to local pO₂. {42. Halpern, 2003 #1798.} The narrow (μT) spectral line-widths, or, equivalently, the inverse transverse relaxation times (1/(5 μs)) of these spin probes are directly proportional to the local oxygen concentration. They give a direct quantitative readout of tissue micro-environmental pO₂. Using spectroscopic EPR imaging {43. Lauterbur, 1984 #177; 44. Maltempo, 1986 #181; 45. Halpern, 1994 #93; 46. Epel, 2008 #2200}, spatial images of quantitative tissue pO₂ may be obtained from tumor and normal tissues of living animals.

Registered Images of Cell Signaling Will Show Local Tissue Response to pO₂ Stimulus.

In some embodiments, the present invention obtains simultaneous registered images of cell signals responding to low pO₂ using nitrogen-centered molecular beacons activated by hypoxia-signaling-coupled reporter proteins. These cell-signal images would be spectrally distinct from the trityl-based pO₂ images and are, in some embodiments, obtained simultaneously with them. At 9 mT, carbon centered and the central manifold of ¹⁴N have sufficiently different absorption frequencies (˜8.4 MHz) that they are, effectively, two-color images. The readout frequencies are low enough to avoid poisonous non-resonant absorption to allow oxygen quantification deep in living tissue. {45. Halpern, 1994 #93.} A different embodiment with simultaneous imaging of pO₂, HIF1α signaling and the vascular endothelial growth factor (VEGF) response to HIF1α would use carbon-centered oxygen-sensitive trityl radicals and ¹⁴N and ¹⁵N molecular beacons, effectively providing three-color images automatically registered with each other.

The present invention, for the first time, provides automatic co-localization of micro-environment stimulus and cell signal response in native animal tissue and tumor environment, allowing their comparison. Distinct responses of normal and tumor tissue will provide insight into therapies that can exploit these differences, targeting malignant tumors and sparing normal tissues. We believe this to be paradigm-shifting work, sharpening the in vivo understanding of signaling process. The present invention will allow monitoring subject-to-subject and tumor-to-tumor variation, allowing a more individualized therapy. This will open a major avenue to the improvement of the therapeutic ratio for cancer therapy.

Section #3: T_(1e) Imaging Using Filtered Backprojection and Single Point Imaging Methodologies

Spin-lattice relaxation (T_(1e)) of a spin probe can be sensitive to various environmental parameters including local oxygen pO₂. We have developed three dimensional pulse imaging of T_(1e) in vivo using fast repetition time saturation, inversion recovery and, stimulated echo sequences. T_(1e) images generated by sequences that have electron spin echo readout are reconstructed with filtered backprojection protocols while those using free induction decay readout are reconstructed with the single point imaging protocols. We compared T_(1e) and T_(2e) imaging of narrow line trityl spin probe in vitro to find their performances very similar. However, for in vivo oxymetry T_(1e) imaging is found to be more promising due to weaker dependence of T_(1e) on environmental factors other than oxygen.

Introduction:

Transverse relaxation or T_(2e) based EPR continuous wave and pulsed imaging techniques have proved to be very promising methodologies for oxygen imaging using injected paramagnetic molecule as the probe of the spatial distribution of oxygen in animals {101. Halpern 1989 #89}; {102. Kuppusamy 1994}; {103. Elas 2003}; {104. Mailer 2006}; {105. Epel 2010}; {106. Subramanian 2002}. In many cases the same relaxation mechanisms that affect T_(2e) of a spin probe, act on spin-lattice relaxation, T_(1e), making T_(1e) sensitive to the environment {107. Slichter, 1996}. For example, spin exchange interaction with molecular oxygen results in linear R_(2e)=1/T_(2e) and R_(1e)=1/T_(1e) dependence on pO₂ with nearly identical proportionality coefficients {108. Ardenkjaer-Larsen, 1998}. On the other hand, not all relaxation mechanisms affect T_(2e) and T_(1e) in the same way. For example, the inter-molecule electron spin dipole interactions will not affect spin-lattice relaxation while enhancing the phase relaxation. This makes T_(1e) imaging a very attractive imaging modality, separate from T_(2e) imaging. However, no systematic attempts to image in vivo T_(1e) using pulsed methods in vivo have been undertaken.

Given the widely used T_(2e) pulsed EPR oxygen imaging techniques, it is natural to use similar techniques to read out T_(1e) based image information. Although very different in imaging principles and observed relaxation kinetics both electron spin echo (ESE) imaging oxymetry and single point imaging (SPI) oxymetry can precisely measure T_(2e). ESE imaging uses filtered backprojection (FBP) for image reconstruction from a number of static gradient projections recorded with fixed gradient amplitude, |{right arrow over (G)}|, and different gradient orientations. The projections are obtained by the Fourier transformation of time domain signals. In order to preserve correct phase information dead-time free time domain signals are required which can be achieved by generation of echoes {104. Mailer, 2006}. In our earlier work we used two pulse π/2-τ-π-τ-echo ESE (2 pESE for brevity) {104. Mailer, 2006}; {109. Epel, 2008 #2200}. To measure T_(2e) multiple separate images with different τ delay values were acquired and the exponential decay times of signal in each individual voxel in those images measured.

SPI methodology is based on different principles {106. Subramanian, 2002}; {110. Matsumoto, 2006}. The dead time free acquisition is achieved by recording the single point on free induction decay, FID, at a known time, t_(SPI), as a function of stepped static gradient amplitudes sampled on a cubic grid. This signal forms a 3D “pseudo-echo”, the FT of which generates a spatial image. The spatial information is encoded into the phase of FID. In the simplest form of SPI, the phase relaxation times are extracted from multiple images obtained at different t_(SPI). The fit of individual voxels to exponential decay gives the FID dephasing time T_(2e)* directly related to T_(2e) as 1/T_(2e)*=1/T_(2e) ^(hf)+1/T_(2e), where T_(2e) ^(hf) is the oxygen independent phase relaxation due to hyperfine interaction with trityl nuclei. More advance imaging schemes were developed to improve the precision and reduce artifacts of SPI technique {110. Matsumoto, 2006}; {111. Devasahayam, 2007}.

For selection of a pulse method for T_(1e) imaging number of pulse sequence parameters should be taken into account. One of them is the pulse sequence bandwidth. Since the whole projection is acquired at once, the bandwidth of the pulse sequence has to be broad enough to cover the equivalent bandwidth of the EPR line broadened by any applied gradient. A bandwidth of about 5-10 MHz is typically sufficient for low frequency in vivo imaging of 3 cm specimens such as portions of mouse anatomy {109. Epel, 2008 #2200}. Another requirement is a dead time free acquisition, which can be achieved by utilizing the same principles implemented in the ESE FBP T_(2e) imaging protocols.

We have selected three conventional ways to determine T_(1e), two of which can be combined with two readout imaging methodologies:

-   -   Saturation by fast repetition (SFR ESE and SFR SPI);     -   Inversion recovery (IRESE and IRSPI);     -   Stimulated echo (SE, ESE only).

In the SFR experiment (FIGS. 1A and 1B) the amplitude of the ESE or FID is measured as a function of repetition time, T_(R), of the respective sequence. A short repetition time (less than T_(1e)) saturates the spins so the echo amplitude reduces as exp(−T_(R)/T_(1e)). Varying the repetition rate monitors the spin system saturation to get T_(1e). The second type of T_(1e) sequence, inversion recover, inverts the spin polarization using a broadband π pulse and the recovery is measured using as a function of the delay T after the inversion pulse. The echo detected inversion recovery sequence with ESE detection, IRESE (FIG. 1A1) has been used for T_(1e) sensitive imaging {112. Eaton, 1987}. The delay τ was kept fixed during experiment. The bandwidth of this sequence is approximately equal to the bandwidth of the 2 pulse ESE detection sequence. The inversion recovery sequence for SPI, referred to as IRSPI, utilizes FID detection (FIG. 1D) and the bandwidth of this sequence is approximately equal to the bandwidth of the inversion pulse. An alternative approach involves saturation recovery experiment. Here, the inversion pulse is replaced by long, t_(p)>>T_(1e), saturation pulse. The bandwidth of this pulse is insufficient for the purposes of imaging. Because additional additional facilities such as modulation of the magnetic field or pulse frequency during this pulse may be required to increase its bandwidth, we excluded this sequence from our evaluation. The third sequence, SE (FIG. 1E), is the three-pulse ‘stimulated’ echo sequence {113. Schweiger, 2001}. In this sequence the π pulse of the 2 pulse ESE experiment is split into two π/2 pulses separated by a waiting time T. After the first two π/2 pulses the magnetization is stored along the z-axis (longitudinal axis) where it remains during the time T. This waiting time is varied allowing the magnetization to decay with T_(1e). The third π/2 pulse rotates the z-component back into the transverse xy-plane where it gives rise to a stimulated echo at fixed time τ after the third pulse. This sequence is known to have nearly double bandwidth as compared to 2 pulse ESE (and thus IRESE) sequence with the same length of the RF pulses {1114. Kevan, 1990} and, therefore, is very promising for reducing applied power in living subject.

We present a study of these different methods for imaging of T_(1e) using our 250 MHz pulse spectrometer. We also compare the precision of T_(1e) imaging with that of T_(2e) imaging for determination of the O₂ concentration in the phantoms.

Materials and Methods:

Spin Probe:

The spin probe used for the EPR imaging was a OX063 radical methyl-tris[8-carboxy-2,2,6,6-tetrakis[2-hydroxyethyl]benzo[1,2-d:4,5-d′]bis[1,3]dithiol-4-yl]-trisodium salt, molecular weight=1,427 from GE Healthcare (Little Chalfont, Buckinghamshire, UK). The 1 mM solution of spin probe in saline was contained in a flat-bottomed borosilicate glass cylinder of 9.5 mm inner diameter and 45 mm length. The 0% O₂ sample was deoxygenated using a multiple-cycles freeze-pump-thaw technique and flame sealed. The 9.3% O₂ sample was produced by bubbling of the solution with corresponding nitrogen-oxygen gas mixture and then was sealed with epoxy. Samples were placed into the resonator horizontally along the resonator's axis of symmetry and centered in the axial plane of the resonator. Because samples were half-full this produced a meniscus at the liquid—air contact surface.

Pulse Imager and Pulse Sequences:

In this work we used the versatile pulse 250 MHz imager described in details elsewhere {109. Epel, 2008 #2200}. To utilize the full power of our 2 kW RF amplifier {115. Quine, 2006} (Tomco Technologies, Norwood S A, Australia) the transmit-receive switch of the imager was redesigned using high power components and utilizing a new protection scheme {116. Sundramoorthy, 2009.} A pulse amplitude modulation switch was added to produce π/2- and π-pulses of equal duration (hence equal bandwidth) {117. Quine, 2010}. The imager control software SpecMan4EPR version 1.1.6 {118. Epel, 2005} was used.

To facilitate the image comparison, the measurement time was kept 10 minutes for all images. Since pulse sequences are intended to image samples with heterogeneous relaxation times, we used the same sequences for phantoms and animal imaging. Standard deviation of relaxation times in a homogeneous phantom was used as an estimation of relaxation time errors. Two outer layers of images were excluded from standard deviation calculations to avoid partial volume averaging artifacts. Tables 1 and 2 present parameters of T_(1e) and T_(2e) sequences. The repetition time for ESE sequences, T_(R), was adjusted to keep the delay between the last pulse in a sequence to the first pulse, T^(LF) _(R), of the next sequence constant. This method of repetition time definition is found to be more efficient as compared to conventional method where repetition time in the experiment is kept constant, independently of sequence length. The optimal repetition time was determined by decreasing T_(R) from 5T_(1e) until sequence still provided correct values for T_(1e) or T_(2e). To accelerate acquisition of the IRESE/IRSPI images, the image corresponding to the last delay T was substituted with an image obtained without inversion pulse and delay T. This halved the acquisition time of the last delay image.

For all ESE sequences the same 3D FBP protocol {104. Mailer, 2006}; {119. Eaton 1991} was applied: 208 projections corresponding to the 18×18 (eighteen-by-eighteen) equal solid angle gradient spacing {120. Ahn 2007(1)} were acquired; gradient strength was |{right arrow over (G)}|=15 mT/m; object field of view was 4.24 cm.

A baseline (acquisition at a far off-resonant field) acquired every fourth trace (53 traces in all). To reduce FBP reconstruction artifacts the acquired set of projections was four-fold linearly interpolated {121. Ahn 2007(2)} and filtered with a 3D Ram-Lak filter with a cutoff at 0.5 times the Nyquist frequency. In the images we kept only those voxels with a signal amplitude greater than 15% of the maximum amplitude at the shortest delay. Further data acquisition and processing methods are discussed in detail elsewhere {109. Epel, 2008 #2200}.

The SPI protocol involved acquisition of 5547 FIDs at delay t_(SPI)=1000 ns with gradients corresponding to 23³ matrix in which only the gradients inside the sphere with the diameter of 23 gradients were taken. The maximum gradient of 15 mT/m was used. A baseline was acquired every 20th trace to suppress imager related artifacts. The 3D ‘pseudo-echo’ matrix was apodized with hamming window and Fourier transformed to produce final image. All data processing was performed using in-house MATLAB (The Mathworks, Inc., and Natick, Mass., USA) software.

Non-Imaging Versus Imaging Conditions:

Acquisition of spatial information requires a considerable time. Therefore, imaging protocols have to balance between precision of spatial and relaxation-time measurements. For example, the relaxation times are estimated from five to eight points on the decay curve. Moreover, the time of in vivo imaging can be limited by the animal physiology. These restrictions do not apply for non imaging measurements on phantoms, which can have large number of delays and can take much longer time. The parameters of sequences for these measurements are selected so that they have no influence on measured relaxation times.

Animal Imaging:

T_(1e) and T_(2e) images were sequentially taken on the same animal with no delay in between. All animal experiments were done according to the USPHS “Policy on Humane Care and Use of Laboratory Animals”, and the protocols were approved by the University of Chicago Institutional Animal Care and Use Committee. The University of Chicago Animal Resources Center is an Association for Assessment and Accreditation of Laboratory Animal Care—approved animal care facility.

Results:

To demonstrate applicability of T_(1e) imaging for oxymetry we present the imaging results on two phantoms with different concentration of O₂: 0% (Table 3, FIG. 2) and 9.3% (Table 4), respectively. This range of oxygen concentrations is important hypoxia studies on animals {103. Elas 2003}; {122. Elas 2008}; {123. Matsumoto 2008}. The relaxation times determined under non-imaging conditions (no applied gradients, T_(R)=15T_(1e)) are given in the footnotes of the tables. The results of T_(2e) measurements using 2 pESE are given in the tables for comparison. The standard deviation of relaxation times in homogeneous phantom was used for estimation of errors. Between all T_(1e) imaging methods IRESE sequence demonstrated the smallest standard deviation. SE showed slightly worse performance, while SFR had very large standard deviations on 0% O₂ phantom and was unable to produce T_(1e) image of 9.3% O₂ phantom. 2 pESE performance superseded the performance of all T_(1e) methods, however on 0% O₂ phantom the performance of IRESE came very close to the performance of 2 pESE. Similarly to ESE, IRSPI sequence demonstrated much better performance than SFR SPI. Here we should note that the comparison of absolute precisions of ESE and SPI methodologies is outside of the scope of this work due to the lack of expertise in SPI.

For demonstration of the T_(1e) imaging on a live animal we selected the IRESE (FIG. 3A) methodology, which showed the best performance on phantoms, and compared the T_(1e) image with T_(2e) image (FIG. 3B) obtained using 2 pESE on the same animal. The outlines and general patterns of the images are very similar. The average relaxation times in T_(2e) image are shorter, consistent with the typical ratios of T_(1e) and T_(2e) in trityls. Nevertheless the breadth of relaxation time histograms is approximately the same.

Discussion:

The data presented in the Results subsection demonstrate the feasibility of T_(1e) imaging of live animals. IRESE images show comparable with T_(2e) images quality both in vitro and in vivo. The major factor that affects precision of T_(1e)/T_(2e) imaging is an image signal to noise ratio (SNR). Assuming that for all methodologies the imager noise characteristics are equal, the image SNR will be governed by an amplitude of a signal and number of acquisitions. Inversion recovery has the highest change in a signal amplitude due to relaxation, double of that for other sequences, since the evolution of signal from negative to positive is monitored. On the other hand the duration of inversion recovery sequences is longer than T_(2e) sequences, which reduces the number of sequence repetitions per unit time and hence SNR. The amplitude of signal for SFR is proportional to exp(−T_(R) ^(MIN)/T_(1e)), where T_(R) ^(MIN) is the minimum T_(R) in an experiment. In our instrument minimum T_(R) is governed by the duty cycle of the power amplifier, insufficient to generate SFR sequences with short enough T_(R). This makes SFR imaging not feasible for our instrument. The performance of SE method is worse than IRESE due to twice lower echo signal {114. Kevan, 1990}. However certain advantages can be derived from considerably smaller power requirements of this sequence. SE image was taken using only 8% of RF power required for other methods excluding SFR SPI (see Table 2). Considering all these factors and nearly equal T_(1e) and T_(2e), inversion recovery methods, IRESE and IRSPI, should have comparable performance to 2 pESE and standard SPI images, which was demonstrated in the experiment.

The reduction in RF power requirements for SE sequence may make it attractive for large subject imaging. The power required for RF pulses with identical B₁ is growing proportional to the resonator volume. Large resonators may require more power than available sources can deliver. The lower average power deposition may also favor this sequence for human applications.

T_(1e) experiments can not deliver correct concentration map of the sample. Being acquired at non-zero τ, the T_(1e) amplitude image will always carry an effect of T_(2e). To obtain true amplitude, the T_(2e) measurement has to be performed and amplitude extrapolated to time 0. Thus it might be of interest to combine inversion recovery T_(1e) and T_(2e) imaging into one experiment. Another advantage of such a joint experiment would be an economy on delays since there will be no need to measure IRESE or IRSPI with long delay T—the result of such measurement is equal to experiment with no inversion pulse (2 pESE or FID SPI).

Conclusions:

T_(1e) imaging is feasible and comparable in precision with T_(2e) imaging. Therefore, it can be considered as an alternative method for oxymetry and other applications. Different applications and instruments may benefit from different T_(1e) methods, with inversion recovery imaging exhibiting the best relaxation time precision while SE imaging having the least power requirements.

FIGS. 4A-4F show various pulse sequences for determination of T_(1e): FIG. 4A shows an SFR ESE (saturation by fast repetition ESE) pulse sequence, wherein the repetition rate is varied, FIG. 4B shows an SFR SPI (saturation by fast repetition SPI) pulse sequence, wherein the repetition rate is varied, FIG. 4C shows an SR (saturation recovery) pulse sequence, wherein the delay time T is varied, FIG. 4D1 shows an IRESE (inversion recovery with ESE detection) pulse sequence, wherein the delay time T is varied, FIG. 4E shows an IRSPI (inversion recovery with SPI detection) pulse sequence, wherein the delay time T is varied, and FIG. 4F shows an SE (stimulated echo) pulse sequence, wherein the delay time T is varied.

Regarding FIG. 5A, FIG. 5B, FIG. 5C, FIG. 5D, FIG. 5E, FIG. 5F: the selected slices and histograms of T1e images are obtained using for FIG. 5A—SFR ESE; for FIG. 5B—SFR SPI; for FIG. 5C—IRESE; for FIG. 5D—IRSPI; for FIG. 5E—SE; and for FIG. 5F—2 pESE. FIG. 5F shows the slice and histogram of T_(2e) 2 pESE image. Experiments A for FIG. 5A, C for FIG. 5C, E for FIG. 5E, and F for FIG. 5F are performed on 1 mM sample 0% O₂, experiments B for FIGS. 5B and D for FIG. 5D are performed on 3 mM 0% O₂ sample of triarylmethyl radical OX063.

FIG. 6A and FIG. 6B show T_(2e) and T_(1e) images obtained using 2 pESE and IRESE pulse sequences, respectively.

Tables

TABLE 1 Pulse sequences. Protocol Description π/2-τ-π-τ-echo; 35 ns π/2 and π RF pulses; τ = 630 ns; 16-step phase cycling, 16640 echoes, including phase cycling; 8 images with different repetition times logarithmically spaced between 10 μs and 25 μs; imaging time 10 minutes. SFR First First Second Second Detection Detection ESE N pulse delay pulse delay channel Re channel Im 1 0.5X 630 ns X 630 ns A B 2 0.5X 630 ns −X 630 ns A B 3 0.5X 630 ns Y 630 ns −A −B 4 0.5X 630 ns −Y 630 ns −A −B 5 −0.5X 630 ns X 630 ns −A −B 6 −0.5X 630 ns −X 630 ns −A −B 7 −0.5X 630 ns Y 630 ns A B 8 −0.5X 630 ns −Y 630 ns A B 9 0.5Y 630 ns X 630 ns B −A 10 0.5Y 630 ns −X 630 ns B −A 11 0.5Y 630 ns Y 630 ns −B A 12 0.5Y 630 ns −Y 630 ns −B A 13 −0.5Y 630 ns X 630 ns −B A 14 −0.5Y 630 ns −X 630 ns −B A 15 −0.5Y 630 ns Y 630 ns B −A 16 −0.5Y 630 ns −Y 630 ns B −A π-T-π/2-τ-π-τ-echo; 35 ns π/2 and π RF pulses; τ = 630 ns; 16-step phase cycling applied only for detection sequence, 7520 acquisitions per T, including phase cycling; 8 T's denoted as VD in the table below logarithmically spaced between 0.5 μs and 16 μs; T^(LF) _(R) = 25 μs; imaging time 10 minutes. First First Second Second Third Third Detection Detection IRESE N pulse delay pulse delay pulse delay channel Re channel Im 1 X VD 0.5X 630 ns X 630 ns A B 2 X VD 0.5X 630 ns −X 630 ns A B 3 X VD 0.5X 630 ns Y 630 ns −A −B 4 X VD 0.5X 630 ns −Y 630 ns −A −B 5 X VD −0.5X 630 ns X 630 ns −A −B 6 X VD −0.5X 630 ns −X 630 ns −A −B 7 X VD −0.5X 630 ns Y 630 ns A B 8 X VD −0.5X 630 ns −Y 630 ns A B 9 X VD 0.5Y 630 ns X 630 ns B −A 10 X VD 0.5Y 630 ns −X 630 ns B −A 11 X VD 0.5Y 630 ns Y 630 ns −B A 12 X VD 0.5Y 630 ns −Y 630 ns −B A 13 X VD −0.5Y 630 ns X 630 ns −B A 14 X VD −0.5Y 630 ns −X 630 ns −B A 15 X VD −0.5Y 630 ns Y 630 ns B −A 16 X VD −0.5Y 630 ns −Y 630 ns B −A π/2-τ-π/2-T-π/2-τ-echo; 60 ns RF pulses; τ = 550 ns; 32-step phase cycling, 12160 acquisitions per T, including phase cycling; 8 T's denoted as VD in the table below logarithmically spaced between 0.45 μs and 7 μs; repetition time 20 μs; imaging time 10 minutes. Detection Detection First First Second Second Third Third channel channel SE N pulse delay pulse delay pulse delay Re Im 1 0.5X 630 ns 0.5X VD 0.5X 630 ns A B 2 0.5X 630 ns 0.5X VD −0.5X 630 ns −A −B 3 0.5X 630 ns 0.5X VD 0.5Y 630 ns −B A 4 0.5X 630 ns 0.5X VD −0.5Y 630 ns B −A 5 0.5X 630 ns −0.5X VD 0.5X 630 ns −A −B 6 0.5X 630 ns −0.5X VD −0.5X 630 ns A B 7 0.5X 630 ns −0.5X VD 0.5Y 630 ns B −A 8 0.5X 630 ns −0.5X VD −0.5Y 630 ns −B A 9 −0.5X 630 ns 0.5X VD 0.5X 630 ns −A −B 10 −0.5X 630 ns 0.5X VD −0.5X 630 ns A B 11 −0.5X 630 ns 0.5X VD 0.5Y 630 ns B −A 12 −0.5X 630 ns 0.5X VD −0.5Y 630 ns −B A 13 −0.5X 630 ns −0.5X VD 0.5X 630 ns A B 14 −0.5X 630 ns −0.5X VD −0.5X 630 ns −A −B 15 −0.5X 630 ns −0.5X VD 0.5Y 630 ns −B A 16 −0.5X 630 ns −0.5X VD −0.5Y 630 ns B −A 17 −0.5Y 630 ns 0.5X VD 0.5X 630 ns B −A 18 −0.5Y 630 ns 0.5X VD −0.5X 630 ns −B A 19 −0.5Y 630 ns 0.5X VD 0.5Y 630 ns A B 20 −0.5Y 630 ns 0.5X VD −0.5Y 630 ns −A −B 21 −0.5Y 630 ns −0.5X VD 0.5X 630 ns −B A 22 −0.5Y 630 ns −0.5X VD −0.5X 630 ns B −A 23 −0.5Y 630 ns −0.5X VD 0.5Y 630 ns −A −B 24 −0.5Y 630 ns −0.5X VD −0.5Y 630 ns A B 25 0.5Y 630 ns 0.5X VD 0.5X 630 ns −B A 26 0.5Y 630 ns 0.5X VD −0.5X 630 ns B −A 27 0.5Y 630 ns 0.5X VD 0.5Y 630 ns −A −B 28 0.5Y 630 ns 0.5X VD −0.5Y 630 ns A B 29 0.5Y 630 ns −0.5X VD 0.5X 630 ns B −A 30 0.5Y 630 ns −0.5X VD −0.5X 630 ns −B A 31 0.5Y 630 ns −0.5X VD 0.5Y 630 ns A B 32 0.5Y 630 ns −0.5X VD −0.5Y 630 ns −A −B π/2-FID; 70 ns π/2 RF pulses; 4-step phase cycling 16640 echoes, including phase cycling; 8 images with different repetition times logarithmically spaced between 10 μs and 25 μs; imaging time 10 minutes. Detection Detection SFR First First channel channel SPI N pulse delay Re Im 1 0.5X 1000 ns A B 2 −0.5X 1000 ns −A −B 3 0.5Y 1000 ns −B A 4 −0.5Y 1000 ns B −A π-T-π/2-did; 35 ns π/2 and π RF pulses; 4-step phase cycling applied only for detection sequence, 7520 acquisitions per T, including phase cycling; 8 T's denoted as VD in the table below logarithmically spaced between 0.5 μs and 16 μs; T^(LF) _(R) = 25 μs; imaging time 10 minutes. Detection Detection First First Second Second channel channel IRSPI N pulse delay pulse delay Re Im 1 X VD 0.5X 1000 ns A B 2 X VD 0.5X 1000 ns −A −B 3 X VD 0.5X 1000 ns −B A 4 X VD 0.5X 1000 ns B −A

TABLE 2 Parameters of pulse sequences Transmitted RF Band- Average Pulse Pulse power width Power Sequence length (W) (MHz) (W) 2pESE (T_(2e)) 35 ns, 39.6(π/2), 8.7 0.56 π/2 and π 158.5(π) SFR ESE (T_(1e)) 35 ns, 39.6(π/2), 8.7 0.4 π/2 and π 158.5(π) IRESE (T_(1e)) 35 ns, 39.6(π/2), 8.7 0.33 π/2 and π 158.5(π) SE (T_(1e)) 60 ns, all π/2 12.6 9.7 0.1

TABLE 3 Relaxation times and precision for 0% O₂ 1 mM sample Standard deviation Pulse T_(2e) or T_(1e) T_(2e) or T_(1e) Sequence (μs) (μs) 2pESE (T_(2e)) 5.15 0.19 SFR ESE (T_(1e)) 6.2 1.4 IRESE (T_(1e)) 5.9 0.29 SE (T_(1e)) 5.8 0.38 Non-imaging relaxation times: T_(2e) = 5.07 μs (2pESE); T_(1e) = 5.8 μs (IRESE); T_(1e) = 5.87 μs (SE).

TABLE 4 Relaxation times and precision for 9.3% O₂ 1 mM sample. Standard deviation Pulse T_(2e) or T_(1e) T_(2e) or T_(1e) Sequence (μs) (μs) 2pESE (T_(2e)) 1.25 0.07 SFR ESE (T_(1e)) N/A* N/A* IRESE(T_(1e)) 1.31 0.15 SE(T_(1e)) 1.35 0.23 *the SFR sequence was unable to produce precise measurement due to T_(1e) considerably smaller than the minimum repetition time that can be achieve in our system. Non-imaging relaxation times: T_(2e) = 1.24 μs (2pESE); T_(1e) = 1.33 μs (IRESE); T_(1e) = 1.31 μs (SE).

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In some embodiments, a series of measurements are made, each using a set of three excitation pulses. In some embodiments, each set of three pulses includes a pulse sequence, wherein each pulse has a phase relative to X-Y polar coordinates, wherein a unity magnitude pulse with zero phase shift is considered to be an 1.0 X pulse (also called simply an X pulse); a unity magnitude pulse with 0.5 pi radian (90 degrees) phase shift is considered to be an 1.0 Y pulse (also called simply a Y pulse); a unity magnitude pulse with 1.0 pi radian (180 degrees) phase shift is considered to be an −1.0 X pulse (also called simply a −X pulse); a unity magnitude pulse with 1.5 pi radians (270 degrees) phase shift is considered to be an −1.0 Y pulse (also called simply a −Y pulse). In like manner, pulses having half that magnitude and a 0-degree phase shift are called 0.5X pulses, those with a 90-degree phase shift are called 0.5Y pulses, those with a 180-degree phase shift are called −0.5X pulses, and those with a 270-degree phase shift are called −0.5Y pulses. In some embodiments, the durations of each of the first, second and third pulse is about 35 ns (about 9 cycles of 250 MHz RF), with a phase delay denoted (+X, +Y, −X, and −Y) and a magnitude of unity (1) or half (0.5). The variable delay between the first pulse and the second pulse is denoted VD1 (also denoted as “t” in FIG. 1B), and the variable delay between the second pulse and the third pulse is denoted VD2 (also denoted as “τ” in FIG. 1B). In some embodiments, a series of sets of pulses are successively generated such that sixteen sets of three pulses are generated for each of one or more first delays VD1 and each of one or more second delays VD2, as follows:

First First Second Second Third Set pulse delay pulse delay pulse 1 −X VD1 +.5X VD2 +Y 2 −X VD1 −.5X VD2 +Y 3 −X VD1 +.5X VD2 −Y 4 −X VD1 −.5X VD2 −Y 5 −X VD1 +.5Y VD2 +X 6 −X VD1 −.5Y VD2 +X 7 −X VD1 +.5Y VD2 −X 8 −X VD1 −.5Y VD2 −X 9 −X VD1 +.5Y VD2 +Y 10 −X VD1 −.5Y VD2 +Y 11 −X VD1 +.5Y VD2 −Y 12 −X VD1 −.5Y VD2 −Y 13 −X VD1 +.5X VD2 +X 14 −X VD1 −.5X VD2 +X 15 −X VD1 +.5X VD2 −X 16 −X VD1 −.5X VD2 −X

In some embodiments, after the third pulse, a third variable delay (also denoted as “τ” in FIG. 1A2) is inserted before a period of the response RF signal is acquired, wherein in some embodiments, the response RF is acquired for a period of about 4 to 6 microseconds. In some embodiments, the magnitude, quadrature (or phase) and/or frequency of the response RF are measured, digitized and stored for later analysis and/or image reconstruction.

In some embodiments, the present invention provides an apparatus for electron paramagnetic resonance imaging (EPRI) of a volume of animal tissue in vivo. This apparatus includes a set of surface transmit coils; and a set of surface receive coils, wherein the set of surface transmit coils generates an excitation magnetic field in the volume of animal tissue in response to an applied electrical signal, and the set of surface receive coils generates a sensed electrical signal in response to a sensed magnetic field in the volume of animal tissue, and wherein the set of transmit coils and the set of receive coils are oriented relative to one another such that the sensed electrical signal has little or no component directly due to the excitation magnetic field, and wherein the set of surface receive coils is configured to detect electron paramagnetic resonance signals in the volume of animal tissue; and a pulsed-RF driver circuit, operatively coupled to the set of transmit coils, that drives a RF pulse set having a plurality of successive RF pulses, including a first pulse having a plurality of cycles of RF, followed by a first delay and thereafter by a second pulse that includes a plurality of cycles of RF that are shifted in phase (by about either zero radians (0 degrees), ½ pi radians (90 degrees), pi radians (180 degrees) or 3/2 pi radians (270 degrees)) relative to the first π-pulse, followed by a second delay and thereafter by a third pulse that includes a plurality of cycles of RF that are shifted in phase (by about either zero radians (0 degrees), ½ pi radians (90 degrees), pi radians (180 degrees) or 3/2 pi radians (270 degrees)) relative to the first pulse. In some embodiments, the animal tissue is human tissue in a living human.

Some embodiments further include a magnetic-field generator configured to generate a substantially static magnetic field in the volume of animal tissue, which is generally orthogonal to the excitation magnetic field and to the sensed magnetic field in the volume of animal tissue, and wherein the excitation magnetic field is generally orthogonal to the sensed magnetic field in the volume of animal tissue; an RF receiver circuit operatively coupled to the set of surface receive coils to receive the sensed electrical signal from the set of surface receive coils and to generate a received electrical signal; a digital-signal processor (DSP) unit operatively coupled to the RF receiver circuit and configured to process the received electrical signal and to generate image data; a storage unit operatively coupled to the DSP unit to receive and store the image data; and a display unit operatively coupled to the storage unit to receive and display the image data.

In some embodiments, the present invention provides a method for electron paramagnetic resonance oxygen imaging (EPROI) of a volume of animal tissue in vivo in an animal. This method includes placing (e.g., by injecting, ingesting, inhaling, swabbing or the like) a reporter molecule in the animal; applying an RF pulse sequence that elicits a T₁ spin-lattice relaxation response from the volume of tissue; and generating an EPRI image of in the animal using the T₁ response.

In some embodiments, the present invention provides a method for electron paramagnetic resonance imaging (EPRI) of a volume of animal tissue in vivo in an animal. This method includes generating a substantially static magnetic field in the volume of animal tissue; generating a excitation set that includes a plurality of RF-excitation magnetic-field pulses including a first pulse that includes a plurality of cycles of RF, followed by a first delay and then a second pulse that includes a plurality of cycles of RF that are shifted in phase (by about either zero radians (0 degrees), ½ pi radians (90 degrees), pi radians (180 degrees) or 3/2 pi radians (270 degrees)) relative to the cycles of the first pulse, followed by a second delay and thereafter a third pulse that includes a plurality of cycles of RF that are shifted in phase (by about either zero radians (0 degrees), ½ pi radians (90 degrees), pi radians (180 degrees) or 3/2 pi radians (270 degrees)) relative to the first pulse in the direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue. The method further includes sensing an RF magnetic field in a direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue, wherein the sensed RF magnetic field is in a direction generally orthogonal to the pulsed excitation magnetic field; generating a received electrical signal based on the sensed RF magnetic field; digitally signal processing the received electrical signal to generate image data; storing the image data; and displaying the image data.

In some embodiments, the present invention provides an apparatus and a corresponding method for improved signal-to-noise (S/N) measurements useful for electron paramagnetic resonance imaging (EPRI), in situ and in vivo, using high-isolation transmit/receive surface coils and temporally spaced pulses of RF energy (e.g., in some embodiments, a first pulse of about 9 cycles of an about-250-MHz signal) having an amplitude sufficient to invert or rotate the magnetization prepared in the temporally static magnetic fields by 180 degrees (a so called pi-pulse (π pulse)) followed closely, but at varied times, by a second radio-frequency pulse of about 9 cycles of substantially the same frequency but having an amplitude half that of the initial pulse to rotate the magnetization by, e.g., 90 degrees (a so called pi-over-two pulse (π/2 pulse)), to the horizontal plane where it evolves for a very short fixed time after which time a third radio-frequency pulse sufficient to rotate the magnetization by, e.g., 180 degrees, that allows the formation of an echo (in some embodiments, the cycles of the second pulse and third pulse are obtained from the same signal source as those of the first pulse but are phase shifted by respect to each other by 0, 90, 180, or 270 degrees to reduce signal artifact), which, in some embodiments, provide improved microenvironmental images that are representative of particular internal structures in the human body and spatially resolved images of tissue/cell protein signals responding to conditions (such as hypoxia) that show the temporal sequence of certain biological processes, and, in some embodiments, that distinguish malignant tissue from healthy tissue.

Embodiments within the scope of the present invention include a computer-readable medium for carrying or having computer-executable instructions or data structures stored thereon. Such computer-readable medium may be any available medium, which is accessible by a general-purpose or special-purpose computer system. By way of example, and not limitation, such computer-readable medium can comprise physical storage medium such as RAM, ROM, EPROM, CD-ROM or other optical-disk storage, magnetic-disk storage or other magnetic-storage devices, EEPROM or FLASH storage devices, or any other medium which can be used to carry or store desired program code means in the form of computer-executable instructions, computer-readable instructions, or data structures and which may be accessed by a general-purpose or special-purpose computer system. This physical storage medium may be fixed to the computer system as in the case of a magnetic drive or removable as in the case of an EEPROM device (e.g., FLASH storage device). In some embodiments, this physical storage medium may be accessible and/or downloadable over the internet.

In some embodiments, the present invention provides an apparatus for electron paramagnetic resonance oxygen imaging (EPROI) of a volume of animal tissue in vivo. This apparatus includes: a set of surface transmit coils; and a set of surface receive coils, wherein the set of surface transmit coils generates an excitation magnetic field in the volume of animal tissue in response to an applied electrical signal, and the set of surface receive coils generates a sensed electrical signal in response to a sensed magnetic field in the volume of animal tissue, and wherein the set of transmit coils and the set of receive coils are oriented relative to one another such that the sensed electrical signal has little or no component directly due to the excitation magnetic field, and wherein the set of surface receive coils is configured to detect electron paramagnetic resonance signals in the volume of animal tissue; a pulsed-RF driver circuit, operatively coupled to the set of transmit coils, that drives a plurality of pulse sets, each pulse set having a plurality of successive transmitted pulses, including a first pulse having a plurality of cycles of RF, followed by a first delay and thereafter by a second pulse that includes a plurality of cycles of RF, followed by a second delay and thereafter by a third pulse that includes a plurality of cycles of RF; and an RF receiver circuit operatively coupled to the set of surface receive coils to receive the sensed electrical signal from the set of surface receive coils and to generate a received electrical signal, wherein the transmitted pulses are of magnitudes and durations configured to measure T₁ spin-lattice relaxation in the volume of tissue.

In some embodiments of the apparatus, for different ones of the plurality of pulse sets, the second transmit pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse, and the third pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse.

In some embodiments of the apparatus, for each of the plurality of pulse sets:

the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians;

the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and

the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians.

In some embodiments of the apparatus, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.

In some embodiments of the apparatus, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, and wherein the cycles of RF have a frequency of about 250 MHz.

In some embodiments of the apparatus, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a t delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values that are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling includes values selected from rows of the following table:

Detec- tion channel Detection First First Second Second Third Third Re channel Im pulse delay pulse delay pulse delay (real) (imaginary) X T 0.5X 630 ns X 630 ns A B X T 0.5X 630 ns −X 630 ns A B X T 0.5X 630 ns Y 630 ns −A −B X T 0.5X 630 ns −Y 630 ns −A −B X T −0.5X 630 ns X 630 ns −A −B X T −0.5X 630 ns −X 630 ns −A −B X T −0.5X 630 ns Y 630 ns A B X T −0.5X 630 ns −Y 630 ns A B X T 0.5Y 630 ns X 630 ns B −A X T 0.5Y 630 ns −X 630 ns B −A X T 0.5Y 630 ns Y 630 ns −B A X T 0.5Y 630 ns −Y 630 ns −B A X T −0.5Y 630 ns X 630 ns −B A X T −0.5Y 630 ns −X 630 ns −B A X T −0.5Y 630 ns Y 630 ns B −A X T −0.5Y 630 ns −Y 630 ns B −A

Some embodiments of the apparatus further include a magnetic-field generator configured to generate a substantially static magnetic field in the volume of animal tissue, which is generally orthogonal to the excitation magnetic field and to the sensed magnetic field in the volume of animal tissue, and wherein the excitation magnetic field is generally orthogonal to the sensed magnetic field in the volume of animal tissue; a digital-signal processor (DSP) unit operatively coupled to the RF receiver circuit and configured to process the received electrical signal and to generate image data; a storage unit operatively coupled to the DSP unit to receive and store the image data; and a display unit operatively coupled to the storage unit to receive and display the image data.

In some embodiments, the present invention provides a method for electron paramagnetic resonance imaging (EPRI) of a volume of animal tissue in vivo in an animal. This method includes generating a substantially static magnetic field in the volume of animal tissue; and generating a plurality of pulse sets, wherein the generating of each one of the plurality of pulse sets includes: generating a first RF excitation magnetic field pulse having a plurality of RF cycles in a first direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue; delaying for a first delay time; generating a second RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field; delaying for a second delay time; generating a third RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field; delaying for a third delay time; sensing an RF spin-relaxation signal; and generating a received electrical signal based on the sensed RF signal; wherein the first, second and third RF excitation magnetic field pulses are of magnitudes and durations configured to measure T₁ spin-lattice relaxation in the volume of tissue. In some embodiments, the animal tissue is human tissue in a living human.

In some embodiments of the method, for different ones of the plurality of pulse sets, the second transmit pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse, and the third pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse.

In some embodiments of the method, for each of the plurality of pulse sets: the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians; the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians.

In some embodiments of the method, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.

In some embodiments of the method, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a t delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, and wherein the cycles of RF have a frequency of about 250 MHz.

In some embodiments of the method, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a t delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values that are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling uses values selected according to the following table:

Detection Detection First First Second Third Third channel channel Im N pulse delay pulse Second delay pulse delay Re (real) (imaginary) 1 X T 0.5X 630 ns X 630 ns A B 2 X T 0.5X 630 ns −X 630 ns A B 3 X T 0.5X 630 ns Y 630 ns −A −B 4 X T 0.5X 630 ns −Y 630 ns −A −B 5 X T −0.5X 630 ns X 630 ns −A −B 6 X T −0.5X 630 ns −X 630 ns −A −B 7 X T −0.5X 630 ns Y 630 ns A B 8 X T −0.5X 630 ns −Y 630 ns A B 9 X T 0.5Y 630 ns X 630 ns B −A 10 X T 0.5Y 630 ns −X 630 ns B −A 11 X T 0.5Y 630 ns Y 630 ns −B A 12 X T 0.5Y 630 ns −Y 630 ns −B A 13 X T −0.5Y 630 ns X 630 ns −B A 14 X T −0.5Y 630 ns −X 630 ns −B A 15 X T −0.5Y 630 ns Y 630 ns B −A 16 X T −0.5Y 630 ns −Y 630 ns B −A

Some embodiments of the method further include digitally signal processing the received electrical signal to generate image data; storing the image data; and displaying the image data.

In some embodiments, the present invention provides an apparatus for electron paramagnetic resonance imaging (EPRI) of a volume of animal tissue in vivo in an animal. This apparatus includes means (as described herein and equivalents thereto) for generating a substantially static magnetic field in the volume of animal tissue; and means for generating a plurality of pulse sets, wherein the means for generating each one of the plurality of pulse sets includes: means for generating a first RF excitation magnetic field pulse having a plurality of RF cycles in a first direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue; means for delaying for a first delay time; means for generating a second RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field; means for sensing an RF spin-relaxation signal; and means for generating a received electrical signal based on the sensed RF signal; wherein the first, second and third RF excitation magnetic field pulses are of magnitudes and durations configured to measure T₁ spin-lattice relaxation in the volume of tissue.

In some embodiments of this apparatus, for different ones of the plurality of pulse sets, the second transmit pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse, and the third pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse. In some embodiments of this apparatus, for each of the plurality of pulse sets: the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians; the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians. In some embodiments of the apparatus, for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.

In some embodiments of the apparatus, for each of the plurality of pulse sets: the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a t delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling uses values selected from rows of the following table:

Detec- tion channel Detection First First Second Second Third Third Re channel Im pulse delay pulse delay pulse delay (real) (imaginary) X T 0.5X 630 ns X 630 ns A B X T 0.5X 630 ns −X 630 ns A B X T 0.5X 630 ns Y 630 ns −A −B X T 0.5X 630 ns −Y 630 ns −A −B X T −0.5X 630 ns X 630 ns −A −B X T −0.5X 630 ns −X 630 ns −A −B X T −0.5X 630 ns Y 630 ns A B X T −0.5X 630 ns −Y 630 ns A B X T 0.5Y 630 ns X 630 ns B −A X T 0.5Y 630 ns −X 630 ns B −A X T 0.5Y 630 ns Y 630 ns −B A X T 0.5Y 630 ns −Y 630 ns −B A X T −0.5Y 630 ns X 630 ns −B A X T −0.5Y 630 ns −X 630 ns −B A X T −0.5Y 630 ns Y 630 ns B −A X T −0.5Y 630 ns −Y 630 ns B −A

Some embodiments of the apparatus further include a magnetic-field generator configured to generate a substantially static magnetic field in the volume of animal tissue, which is generally orthogonal to the excitation magnetic field and to the sensed magnetic field in the volume of animal tissue, and wherein the excitation magnetic field is generally orthogonal to the sensed magnetic field in the volume of animal tissue; a digital-signal processor (DSP) unit operatively coupled to the RF receiver circuit and configured to process the received electrical signal and to generate image data; a storage unit operatively coupled to the DSP unit to receive and store the image data; and a display unit operatively coupled to the storage unit to receive and display the image data.

It is to be understood that the above description is intended to be illustrative, and not restrictive. Although numerous characteristics and advantages of various embodiments as described herein have been set forth in the foregoing description, together with details of the structure and function of various embodiments, many other embodiments and changes to details will be apparent to those of skill in the art upon reviewing the above description. The scope of the invention should, therefore, be determined with reference to the appended claims, along with the full scope of equivalents to which such claims are entitled. In the appended claims, the terms “including” and “in which” are used as the plain-English equivalents of the respective terms “comprising” and “wherein,” respectively. Moreover, the terms “first,” “second,” and “third,” etc., are used merely as labels, and are not intended to impose numerical requirements on their objects. 

What is claimed is:
 1. A method for electron paramagnetic resonance oxygen imaging (EPROI) of a volume of animal tissue in vivo in an animal, the method comprising: placing a reporter molecule in the animal; establishing a static (DC) magnetic field on the volume of animal tissue; applying an RF pulse sequence that elicits a T_(1e) spin-lattice relaxation response from the volume of tissue; and generating an EPROI image of oxygen in the animal using the T_(1e) response.
 2. The method of claim 1, the applying of the RF pulse sequence further comprising: generating a plurality of pulse sets, wherein the generating of each one of the plurality of pulse sets includes: generating a first RF excitation magnetic field pulse having a plurality of RF cycles in a first direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue, delaying for a first delay time, generating a second RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field, delaying for a second delay time, generating a third RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field, delaying for a third delay time, sensing an RF spin-relaxation signal, and generating a received electrical signal based on the sensed RF signal; wherein the first, second and third RF excitation magnetic field pulses are of magnitudes and durations configured to measure T_(1e) spin-lattice relaxation in the volume of tissue.
 3. The method of claim 1, wherein for each of the plurality of pulse sets: the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians; the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians.
 4. The method of claim 1, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.
 5. The method of claim 1, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, and wherein the cycles of RF have a frequency of about 250 MHz.
 6. The method of claim 1, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values that are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling uses values selected according to the following table: Detection Detection First First Second Second Third channel channel Im N pulse delay pulse delay Third pulse delay Re (real) (imaginary) 1 X T 0.5X 630 ns X 630 ns A B 2 X T 0.5X 630 ns −X 630 ns A B 3 X T 0.5X 630 ns Y 630 ns −A −B 4 X T 0.5X 630 ns −Y 630 ns −A −B 5 X T −0.5X 630 ns X 630 ns −A −B 6 X T −0.5X 630 ns −X 630 ns −A −B 7 X T −0.5X 630 ns Y 630 ns A B 8 X T −0.5X 630 ns −Y 630 ns A B 9 X T 0.5Y 630 ns X 630 ns B −A 10 X T 0.5Y 630 ns −X 630 ns B −A 11 X T 0.5Y 630 ns Y 630 ns −B A 12 X T 0.5Y 630 ns −Y 630 ns −B A 13 X T −0.5Y 630 ns X 630 ns −B A 14 X T −0.5Y 630 ns −X 630 ns −B A 15 X T −0.5Y 630 ns Y 630 ns B −A 16 X T −0.5Y 630 ns −Y 630 ns B −A


7. The method of claim 1, further comprising: digitally signal processing the received electrical signal to generate image data; storing the image data; and displaying the image data.
 8. The method of claim 1, wherein the animal tissue is human tissue in a living human.
 9. An apparatus for electron paramagnetic resonance imaging (EPRI) of a volume of animal tissue in vivo in an animal, the apparatus comprising: means for generating a substantially static magnetic field in the volume of animal tissue; and means for generating a plurality of pulse sets, wherein the means for generating each one of the plurality of pulse sets includes: means for generating a first RF excitation magnetic field pulse having a plurality of RF cycles in a first direction generally orthogonal to the substantially static magnetic field in the volume of animal tissue from a surface of the animal next to the volume of animal tissue; means for delaying for a first delay time; means for generating a second RF excitation magnetic field pulse having a plurality of RF in the first direction generally orthogonal to the substantially static magnetic field; means for sensing an RF spin-relaxation signal; and means for generating a received electrical signal based on the sensed RF signal; wherein the first, second and third RF excitation magnetic field pulses are of magnitudes and durations configured to measure T_(1e) spin-lattice relaxation response in the volume of tissue; and means for generating an EPROI image of oxygen in the animal using the T_(1e) response.
 10. The apparatus of claim 9, wherein for different ones of the plurality of pulse sets, the second transmit pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees for different ones of the plurality of pulse sets) relative to the first pulse, and the third pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse.
 11. The apparatus of claim 9, wherein for each of the plurality of pulse sets: the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians; the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians.
 12. The apparatus of claim 9, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.
 13. The apparatus of claim 9, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling uses values selected from rows of the following table: Detec- tion channel Detection First First Second Second Third Third Re channel Im pulse delay pulse delay pulse delay (real) (imaginary) X T 0.5X 630 ns X 630 ns A B X T 0.5X 630 ns −X 630 ns A B X T 0.5X 630 ns Y 630 ns −A −B X T 0.5X 630 ns −Y 630 ns −A −B X T −0.5X 630 ns X 630 ns −A −B X T −0.5X 630 ns −X 630 ns −A −B X T −0.5X 630 ns Y 630 ns A B X T −0.5X 630 ns −Y 630 ns A B X T 0.5Y 630 ns X 630 ns B −A X T 0.5Y 630 ns −X 630 ns B −A X T 0.5Y 630 ns Y 630 ns −B A X T 0.5Y 630 ns −Y 630 ns −B A X T −0.5Y 630 ns X 630 ns −B A X T −0.5Y 630 ns −X 630 ns −B A X T −0.5Y 630 ns Y 630 ns B −A X T −0.5Y 630 ns −Y 630 ns B −A


14. An apparatus for electron paramagnetic resonance oxygen imaging (EPROI) of a volume of animal tissue in vivo, the apparatus comprising: a set of transmit coils; and a set of receive coils, wherein the set of surface transmit coils generates an excitation magnetic field in the volume of animal tissue in response to an applied electrical signal, and the set of surface receive coils generates a sensed electrical signal in response to a sensed magnetic field in the volume of animal tissue, and wherein the set of surface receive coils is configured to detect electron paramagnetic resonance signals in the volume of animal tissue; a pulsed-RF driver circuit, operatively coupled to the set of transmit coils, that drives a plurality of pulse sets, each pulse set having a plurality of successive transmitted pulses, including a first pulse having a plurality of cycles of RF, followed by a first delay and thereafter by a second pulse that includes a plurality of cycles of RF, followed by a second delay and thereafter by a third pulse that includes a plurality of cycles of RF; and an RF receiver circuit operatively coupled to the set of surface receive coils to receive the sensed electrical signal from the set of surface receive coils and to generate a received electrical signal, wherein the transmitted pulses are of magnitudes and durations configured to measure an electron T_(1e) spin-lattice relaxation in the volume of tissue in order to measure oxygen in the volume of tissue.
 15. The apparatus of claim 14, wherein for different ones of the plurality of pulse sets, the second transmit pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse, and the third pulses have RF cycles that are shifted in phase by a selected different amount (by about either 0 degrees, 90 degrees, 180 degrees or 270 degrees) relative to the first pulse.
 16. The apparatus of claim 14, wherein for each of the plurality of pulse sets: the first transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin by pi radians; the second transmit pulse is a pi/2 pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin ½ pi radians; and the third transmit pulse is a pi pulse having a magnitude and duration selected to rotate an electron paramagnetic resonance spin pi radians.
 17. The apparatus of claim 14, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence.
 18. The apparatus of claim 14, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a t delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, and wherein the cycles of RF have a frequency of about 250 MHz.
 19. The apparatus of claim 14, wherein for each of the plurality of pulse sets, the first, second and third transmit pulses form an inversion recovery with electron-spin echo detection (IRESE) sequence having a π-pulse as the first pulse, a T delay as the first delay, a π/2-pulse as the second pulse, a τ delay as the second delay, a π-pulse as the third pulse, a τ delay as the third delay, wherein the first, second and third pulses are each about 35 ns in duration, wherein the π/2 pulse rotates a magnetization π/2 radians and the π-pulses rotate a magnetization π radians, wherein τ=630 ns, wherein T has a value in a range of about 500 ns to about 16,000 ns, wherein the cycles of RF have a frequency of about 250 MHz, wherein the plurality of pulse sets apply a sixteen-step phase cycling, wherein about 7520 acquisitions are acquired per value of T and include phase cycling, wherein eight T values that are approximately logarithmically spaced between one-half microseconds (0.5 μs) and sixteen microseconds (16 μs) are used, wherein T^(LF) _(R)=25 μs, and wherein the phase cycling includes values selected from rows of the following table: Detec- tion channel Detection First First Second Second Third Third Re channel Im pulse delay pulse delay pulse delay (real) (imaginary) X T 0.5X 630 ns X 630 ns A B X T 0.5X 630 ns −X 630 ns A B X T 0.5X 630 ns Y 630 ns −A −B X T 0.5X 630 ns −Y 630 ns −A −B X T −0.5X 630 ns X 630 ns −A −B X T −0.5X 630 ns −X 630 ns −A −B X T −0.5X 630 ns Y 630 ns A B X T −0.5X 630 ns −Y 630 ns A B X T 0.5Y 630 ns X 630 ns B −A X T 0.5Y 630 ns −X 630 ns B −A X T 0.5Y 630 ns Y 630 ns −B A X T 0.5Y 630 ns −Y 630 ns −B A X T −0.5Y 630 ns X 630 ns −B A X T −0.5Y 630 ns −X 630 ns −B A X T −0.5Y 630 ns Y 630 ns B −A X T −0.5Y 630 ns −Y 630 ns B −A


20. The apparatus of claim 14, further comprising: a magnetic-field generator configured to generate a substantially static magnetic field in the volume of animal tissue, which is generally orthogonal to the excitation magnetic field and to the sensed magnetic field in the volume of animal tissue, and wherein the excitation magnetic field is generally orthogonal to the sensed magnetic field in the volume of animal tissue; a digital-signal processor (DSP) unit operatively coupled to the RF receiver circuit and configured to process the received electrical signal and to generate image data; a storage unit operatively coupled to the DSP unit to receive and store the image data; and a display unit operatively coupled to the storage unit to receive and display the image data. 